Biologically integrated electrode devices

ABSTRACT

Bioelectrodes having enhanced biocompatible and biomimetic features are provided. Methods of making and using the bioelectrodes are further provided. A biologically integrated bioelectrode device and method for detecting electronic signals using a bioelectrode comprising a first electrically conductive substrate and a biological component. The bioelectrode also comprises a conductive polymer electrically coupling the first electrically conductive substrate and the biological component to define a bioelectrode. The bioelectrode can transmit or receive an electrical signal between the electrically conductive substrate and the biological component and conductive polymer.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional Application No.60/713,070, filed on Aug. 31, 2005. The disclosure of the aboveapplication is incorporated herein by reference.

GOVERNMENT RIGHTS

This disclosure was made with government support under National ScienceFoundation Grant No. DMR-0084304 and National Institutes of Health GrantNo. N01-NS-1-2338. The Government has certain rights in the invention.

FIELD

The present teachings relates to biocompatible, biologically integratedbioelectrode devices resulting from non-toxic deposition andpolymerization of conducting polymers, in the presence of biologicalcomponents. In particular, the present teachings relates to an apparatusand methods for the detection, stimulation, and recording of electrical,chemical, and ionic interactions between a bioelectrode and variousbiologic and chemical targets. The methods can be used for the detectionand stimulation of charge transfer interactions between a conductivepolymer and the surrounding tissue, cells, chemicals, electrolytes,charge carriers receptors and enzymes that are permitted to interactwith the bioelectrode.

BACKGROUND

Inherently “conductive polymers” (π-conjugated conductive polymers) andnon-conductive polymers with conductive dopants are useful asbiocompatible polymeric coating materials for preexisting electrodes,probes, and sensors providing unique electrical, biochemical andelectroactive properties. The monomers that polymerize to formconductive polymers can comprise one or more of3,4-ethylenedioxythiophene (EDOT), pyrrole, anilines, acetylenes,thiophenes, and blends thereof.

Surface and bulk materials currently used as electrodes for biomedicaldevices offer limited biocompatibility, resulting in tissue injury andinflammation in the vicinity of the implanted device. In addition tolimited biocompatibility, stimulation of chronic negative immune systemreactions often lead to biofouling of existing implants of electrodesand erosion of device surface materials. Various biological tissues,including the central nervous system (CNS) react negatively to implanteddevices, varying in severity according to the site of implantation, thematerials used and differences in electrode geometries and implantationmethodologies. Chronic rejection of the implantable devices in the CNScan be characterized by a hypertrophic reaction from surroundingastrocytes with increased expression of filament proteins and vimentin.In addition to protein adsorption to the device surface, microglialcells and foreign body giant cells envelop the implanted devicesresulting in encapsulation of the device and formation of highelectrical impedance fibrous scar tissue. This diminishes, andeventually negates signal transduction between the tissue and thedevice. Similar foreign-body responses are found throughout human andanimal tissues including major targets for novel implanted biomedicaldevices including the brain, heart, and skin. Bioincompatibilityrepresents a key weakness of new implantable biomedical devicescurrently being developed and is the foremost roadblock to successful invivo testing and usage.

Surface modification of implantable electrodes and sensors shouldprovide improvements in both their long term biocompatibility andeletro-functionality. It would be highly desirable to design electrodedevices which could intimately interface electrode sites to livingtissue, as well as to facilitate efficient charge transport fromionically conductive tissue to the electronically conductive electrodeand induce surrounding tissue to attach or interface directly to theimplanted device.

SUMMARY

The present teachings provide a composition for the construction of anelectrochemical sensing and stimulation device wherein the electrode isintimately in contact with a biological component during the recordationand stimulating process. An electrically conductive substrate or workingelectrode is coated with a conductive polymer that can be polymerized asa conductive polymer matrix in the presence of live tissue, cells, cellconstituents and in artificial scaffolds which greatly increases theeffective surface area of the bioelectrode resulting in loweredimpedance and enhanced biocompatibility thereby facilitating signaltransduction. The biologically integrated electrode further stabilizesthe electrically conductive substrate by interfacing the electricallyconductive substrate with surrounding cells and/or tissue when implantedand can be loaded with bioactive substances that prevent the formationof unwanted immune rejections.

A further aspect of the present teachings relates to biologicallyintegrated electrodes comprising a first electrically conductivesubstrate and a biological component. The bioelectrode also comprises aconductive polymer comprising one or more conductive polymerselectrically coupling the first conductive substrate to the biologicalcomponent to define a bioelectrode. The bioelectrode can transmit orreceive an electronic signal between the first electrically conductivesubstrate and the biological component and conductive polymer.

A further aspect of the present teachings relates to a method ofelectrically detecting the transfer of charge between or within cells inliving tissue. The method includes the steps of providing a bioelectrodedevice comprising a first electrically conductive substrate or workingelectrode in intimate contact with tissue capable of transferringelectronic charge. The bioelectrode device is made up of a firstelectrically conductive substrate and a biological component. Thebioelectrode also includes a conductive polymer electrically couplingthe first electrically conductive substrate to the biological componentto collectively define a bioelectrode. The bioelectrode can transmit orreceive an electrical signal between the first electrically conductivesubstrate any one of the biological component and the conductivepolymer. The method also includes electrically connecting thebioelectrode device and a second electrically conductive substrate(another electrode) electrically coupled with the bioelectrode to apower source. The method further includes applying a voltage or currentacross the first and second electrically conductive substrates, therebyinducing a voltage or current across the conductive polymer. The methodalso detects the transfer of electrical signals with the bioelectrodedevice.

Further areas of applicability of the present teachings will becomeapparent from the detailed description provided hereinafter. It shouldbe understood that the detailed description and specific examples, whileindicating certain embodiments of the present teachings, are intendedfor purposes of illustration only and are not intended to limit thescope of the invention.

DRAWINGS

The drawings described herein are for illustration purposes only and arenot intended to limit the scope of the present disclosure in any way.

FIGS. 1A and 1B depict illustrations of a cell adhered to a conductivesubstrate and contacted with a conducting polymer in accordance with theteachings of the present disclosure. FIG.1A is an illustration depictingan adherent cell on an electrically conductive substrate surface insolution with conducting monomer and corresponding counter ions ordopant prior to application of current and electropolymerization of theconducting monomer into conducting polymer around the substrate andportions of the adherent cell. FIG. 1B is an illustration depicting anadherent cell on an electrically conductive substrate with electricallyconductive polymerization of the monomer around the substrate andportions of the adherent cell in accordance with the present disclosure.

FIGS. 2A-2C depict photomicrographs of scanning electron micrographs ofneurons embedded within electropolymerized conducting polymer. FIG. 2Adepicts an electron photomicrograph showing a neuron integratedcompletely and intimately with PEDOT. FIG. 2B is a magnification of FIG.2A depicting three-dimensional fuzzy electrically conductive polymeraround the attached neuron. FIG. 2C depicts a higher magnification ofFIG. 2B showing intimate contact of the electrically conductive polymerin contact with neuron processes according to the teachings of thepresent disclosure.

FIGS. 3A-3C depict photomicrographs of electron microscopic examinationof templated electrically conductive polymer forming cell-definedpolymer topography including cavities and crevices suitable for cellularoccupation. FIG. 3A depicts neuron templated tunnels and neuriteinvaginations on the surface of an electrically conductive substrate.FIG. 3B depicts a magnified neurite outgrowth templated tunnel. FIG. 3Cdepicts a closer magnification of the conducting polymer templatedtunnel according to the teachings of the present disclosure.

FIGS. 4A-4C are line graphs illustrating the electrochemicalcharacteristics of bare electrodes, electrodes coated with conductivepolymer, cell templated conductive polymer and electrodes with adherentcells surrounded by conductive polymer. FIG. 4A depicts the reduction ofelectrical impedance in electrodes coated with conductive polymer andcells templated and embedded with conductive polymer over bareelectrodes. FIG. 4B depicts phase angle reduction in electrodes coatedwith conductive polymer and cells templated and embedded with conductivepolymer over bare electrodes. FIG. 4C depicts normalized impedance overcharge density of electrodes coated with conductive polymer incomparison to templated polymer containing and cell embedded electrodes.

FIG. 5 depicts an illustration of a space-filling electrode according tothe teachings of the present disclosure.

FIGS. 6A and 6B depict line graphs of electrode electrochemicalcharacteristics when implanted in hydrogel constructs in cochlea of aliving guinea pig in accordance with the present disclosure. FIG. 6Adepicts a line graph of electrical impedance over frequency. FIG. 6Bdepicts a line graph of cyclic voltammetry analysis of electrodesembedded into cochlea of a living guinea pig in accordance with theteachings of the present disclosure.

FIG. 7 depicts an illustration of components used to fabricate an insitu bioelectrode device in tissue in accordance with the teachings ofthe present disclosure.

DETAILED DESCRIPTION

The biocompatible electrodes (bioelectrodes), modified electrodes andcoatings contemplated by certain embodiments of the present teachingsinclude electrode devices and/or conductive polymer coatings which havelow biodegradability, low electrical impedance, long-term electricalstability under in vivo conditions, are mechanically soft, are highlybiomimetic (cell feature/cell surface templated & patterned) withnanometer and micrometer scale surface features. Certain embodiments ofthe present teachings relate to conducting cell-templated, livecell-seeded bioelectrodes, or molecular electrode networks that can beadapted for any molecular species including, for example: capable offorming conductive macromolecular networks; biocompatible non-toxicand/or non-immunogenic in the polymer or macromolecular state; able tobe processed from a water or saline-based solution or gel of monomericunits or small oligomer components into macromolecular or polymericfilms or networks by electrochemical or chemical polymerization orcrosslinking with either UV/photo, electrical, thermal, chemical, orself-initiation. The resulting electroconductive films and electrodecoatings can be mechanically stable to withstand degradation andmaintain electrical integrity and connectivity for the duration ofimplementation.

Devices having macro, micro and nano-scale components can be patternedwith such electroconductive polymers polymerized in the presence ofbiological components including tissue, cells, cellular constituentsincluding, membranes, receptors, antibodies, ion-channels, growthfactors and other biological molecules and agents. Bioelectrodes,modified electrodes and electrode coatings of the present teachingsimpart beneficial features including electrodes and electrode coatingmaterials that are electrically stable over time following implantationin tissue, relatively non-biodegradable yet biocompatible, elicitinglower levels of immuno-reactivity than commonly used conductivesubstrate materials such as silicon, platinum, iridium, indium tinoxide, and tungsten. The bioelectrodes or electrode coatings of thepresent teachings can be readily modified to contain a variety ofbioactive agents to facilitate interactions with specific proteins orbiomolecules on the target cells and can limit non-specific interactionsthat are associated with device surface biofouling. Proteins can beincorporated into the conducting polymer material via a variety ofmethods such as electrochemical deposition, covalent linkage, andentrapment in the polymer matrix. Bioelectrodes and devices comprisingelectrode coatings described herein, can be soft, fuzzy electrodes withlow electrical impedance and large surface areas with biomimetic surfacepatterns (cell-shaped holes and tunnels with cell-surface templatednanoscale features). The large surface area is ideal for facilitatingmaximal charge transfer between the electrode and target tissue. Thepliability of the polymer allows for decreased mechanical strain at theinterface between the soft tissue and the hard device surface comparedto a rigid metal electrode. Together, these qualities allow theconductive polymer-coated cells, cell components, and bioactivemolecules of the present teachings to serve as a high surface area,soft, biocompatible, and electrically stable surface coatings forexisting electrode-based biomedical devices that will result indecreased immunoreactivity and improved signal transduction andintegration (tissue adhesion) at the interface between the tissue andthe device.

In some embodiments, the uses of conducting polymers patterned on thesurface of electrically conductive substrates facilitate signaltransport from onically conductive tissue to the electronicallyconductive electrode. Polymerized conducting polymer with a biologicalcomponent is also referred to herein as a “conducting polymer network”or “hybrid biological component-conducting polymer material” to describethe three dimensional nature of the conductive substrate coating.Certain embodiments of the present teachings provide for novelconducting polymer networks as well as a process for polymerizingconducting polymers in the presence of tissue, cells, cell constituentsand other bioactive molecules that result in intimate, directinterfacing between the surface of an electrode-based device and abiological environment.

The detailed description of the present teachings will deal separatelywith the electrode components, including, electrically conductivesubstrates, conjugated conducting polymers, biological components,optional instrumentation including controllers and analyticalinstruments and power sources. Methods of fabricating the variousexemplified bioelectrodes and electrodes modified with the biologicalcomponent embedded conducting polymer coatings and their uses arefurther described. Finally, the present teachings are exemplified with anumber of bioelectrodes and devices and experiments demonstrating theutility and novelty thereof.

A. Device Components and Materials

I. Electrically Conductive Substrates

Electrode substrates can comprise any conducting material or combinationof conducting and non-conducting materials. A number of exemplaryelectrically conductive substrate configurations are described and canbe understood that other configurations can be used. In non-limitingembodiments, electrically conductive substrates can be manufactured frommetals including, but not limited to: Gold (Au), Platinum (Pt), Iridium(Ir), Palladium (Pd), Tungsten (W), Nickel (Ni), Copper (Cu) Aluminum(Al), Stainless Steel (SS), Indium-Tin-Oxide (ITO), Zinc (Zn), Titanium(Ti), Tungsten (W) and their alloys and oxides. Other electricallyconductive substrates can include: carbon, carbon fiber, glassy carbon,carbon composites, carbon paste, conductive ceramics, for example, dopedsilicon (Si), conductive monomers and polymers. As used herein, thefirst electrically conductive substrate is the substrate or electrodethat is in contact or coupled with the biological component and theconducting polymer. The first electrically conductive substrate can alsobe referred to as the working electrode in a multi conductive substratebioelectrode device. The second electrically conductive substrate canalso be referred to any one of: the reference electrode, the counterelectrode, or the saturated calomel electrode.

In some embodiments, the electrode can be patterned with electricallyconducting material such as metal powders, conductive polymers orconductive ceramics. The underlying support material need notnecessarily be composed of conducting material, provided that thesupport material can be made conductive, or that conductive material canbe formed or patterned in or on the non-conductive support.

Devices comprising one or more electrode arrays can include any suitablesupport material upon which a plurality of conducting material channels,dots, spots are formed. In general, if the support material of theelectrode is to come into contact with biological fluid, the supportshould be essentially biocompatible. The microelectrode arrays of thepresent teachings need not be in any specific shape, that is, theelectrodes need not be in a square matrix shape. Contemplated electrodearray geometries can include: squares; rectangles; rectilinear andhexagonal grid arrays various polygon boundaries; concentric circle gridgeometries wherein the electrodes form concentric circles about a commoncenter, and which may be bounded by an arbitrary polygon; and fractalgrid array geometries having electrodes with the same or differentdiameters. Interlaced electrodes can also be used in accordance with thepresent teachings. In some embodiments, the array of electrodes cancomprise about 9 to about 16 electrodes in a 4×4 matrix, 16 to about 25electrodes in about a 5×5 matrix, 10 to 100 electrodes in a 10×10matrix. Other sized arrays known in the art may be used in accordancewith the present teachings.

Production of patterned array of microelectrodes can be achieved by avariety of microprinting methodologies commonly known in the productionof micro-arrays, including, without limitation, by ejecting a pluralityof electro-conducting polymers via a multi-line head nozzle, viaink-jetting techniques and the like. They can be patterned usingphotolithographic and etching methods known for computer chipmanufacture. The micromechanical components may be fabricated usingcompatible “micromachining” processes that selectively etch away partsof the silicon wafer, or comparable substrate, or add new structurallayers to form the mechanical and/or electromechanical components.

Electrodes formed on polymeric supports such as those contemplated inMicro-electro-mechanical systems (MEMS) manufacture can includedepositing thin films of conducting material on a support material,applying a patterned mask on top of the films by photolithographicimaging or other known lithographic methods, and selectively etching thefilms. A thin film may have a thickness in the range of a few nanometersto 100 micrometers. Deposition of electroconducting materials for use asmicro or nano electrodes contemplated in the present teachings can alsoinclude chemical procedures such as chemical vapor deposition (CVD),electrodeposition, epitaxy and thermal oxidation and physical procedureslike physical vapor deposition (PVD) and casting. Methods formanufacture of nano-electromechanical systems having enhancedbiocompatible interfaces comprising conducting polymers andbiomolecules, including cells and cell constituents may be used forcertain embodiments of the present disclosure. (See, for example,Craighead, Science 290: 1532-36, 2000).

In some embodiments of the present teachings, an array or subarrays ofconducting polymer comprising one or more cells, cell constituentsand/or bioactive molecules on an electrode can be connected to variousfluid filled compartments, (including conducting monomer solutionscomprising cells, cell constituents, hydrogel and biological molecules),such as microfluidic channels, nanochannels and/or microchannels. Theseand other components of the apparatus may be formed as a single unit,for example in the form of a chip, microcapillary or microfluidic chips.Various forms of microfabricated chips may be commercially availablefrom, for example, Caliper Technologies Inc. (Mountain View, Calif.,USA.) and ACLARA BioSciences Inc. (Mountain View, Calif., USA).

In some embodiments, the degradability of the electrode substrate can bedependent on the function served by the device or electrode. Forexample, conductive and non-conductive materials with conductivedegradable polymers can be synthesized out of materials including PLGA,PLA, HA, biorubber, oxide glass and other biocompatible biodegradablematerials known to those skilled in the art. The relevance of suchmaterials is apparent when the function of the device or electrode is totransiently stimulate, regenerate injured or defective tissue and thenfade from prevalence after successful implantation to make room forcomplete regeneration and connectivity of the cells or tissue. In someembodiments, the electrically conductive substrate can be permanent tosemi-permanent, wherein the device may be used for extended periods oftime or once implanted, it would be deleterious to remove, for examplesome deep brain neural prosthesis, heart pacemakers and the like.Electrically conductive substrates contemplated for such long-term usagecan include metals, ceramics, and non-degradable conducting polymers,for example, PEDOT.

In some embodiments, the electrode can be connected in part or in wholeto other device components, including wires, leads, conductive polymersthat are in electrical communication with other device components usedto measure, record and analyze the flow of current or detect changes inimpedance, inductance, resistance or capacitance of the bioelectrode,cell, conducting polymer-cell interface or site of implantation. Invarious embodiments of the present teachings multiple or a plurality ofelectrodes in parallel or in series can be used to polymerize theconducting monomer, perform electrochemical oxidation/reductionreactions, provide a current or currents and voltages to stimulatetissue and/or cells, for release of bioactives and for recording orsensing electrochemical events. Other electrodes that can be implementedin the devices described herein can further include various counterelectrodes and saturated calomel electrodes or reference electrodes.

II. Conductive Polymer

In certain embodiments of the present teachings, conductive polymers canimpart desirable features, for example: are electrically stable overtime following implantation in tissue, relatively non-biodegradable yethighly biocompatible, eliciting lower levels of immunoreactivity thancommonly used conducting materials such as silicon, platinum, iridium,indium tin oxide, and tungsten. As used herein, conductive polymers areconjugated polymers that are capable of conducting electrons. The term“conductive polymer(s)” is used interchangeably with “conductingpolymer(s)”. Conductive polymers are formed from their monomeric formvia electrochemical polymerization, oxidative polymerization and othermethods commonly used in the art. Conducting polymer polymerized aroundan electrically conductive substrate can also be referred to as aconducting polymer network due to its three dimensional, fuzzy, softfibrils that extend out from the electrically conductive substrate. Insome embodiments, the conducting polymer network contains embeddedbiological components including cells, cellular constituents, bioactivemolecules or substances and combinations thereof. Conducting polymernetworks having one or more biological components are also referred toas hybrid biological component-conducting polymer material. In certainembodiments of the present teachings, the conductive polymers can bepolymerized in the presence of dopants, tissue, cells, cell parts,cellular constituents, other bioactive molecules, viral, plasmid, yeast,dendromer, quantum dot, or micro-nano particle gene delivery vectors,and/or biodegradable micro-nano particles or fibers that are comprisedof naturally-derived or synthetic polymers that may be decorated withsurface functional groups or molecules intended for interaction withspecific cells or molecules in the target tissue or may be employed forcontrolled-release delivery of bioactive molecules contained within.

In some embodiments, the conducting polymers can include, but are notlimited to: poly(3,4-ethylenedioxythiophene) (PEDOT), poly(pyrrole),polyanilines, polyacetylenes, poly (diallyldimethylammonium chloride,poly-4-vinylpyridine, poly(vinylalcohol), polythiophenes, polymer blendsthereof, and composites with the ability to conduct electronic charge orions, and hybrid polymer-metal materials that are electrically orionically conductive. Other conductive polymer can includefunctionalized copolymers made from EDOT and other conducting polymerderivatives, functional groups such as RGD, IKVAV, YIGSR peptides, andother functional groups that can be covalently attached to theconducting monomer, or they can be linked to spacers having bifunctionalmoities that can be attach to the conducting monomer. A covalentattachment can be effected using any covalent chemistry known in theart. Examples of preferred covalent attachment chemistries includeamine, amide, ester, ether, and their heteroatom cognates, e.g.,sulfonamide, thioether, and so forth. Typically, each pair of entitiesto be joined can jointly comprise a pair of reactive groups, such as anucleophile and an electrophile, one respectively on each member of thepair. Where the biological entity (biomolecule, cell, cell fragment,organelle, or other biologic) is to be directly attached to the monomeror polymer, each will contain one reactive group of a pair. Whereattachment is to take place through a linker, the linker will containtwo reactive groups, one of which is capable of covalently reacting witha reactive group of the monomer and the other of which is capable ofcovalently reacting with a reactive group of the biological entity. Thereactive group(s) can be already present as part of the monomer, linker,or biological entity, or it can be added thereto by reaction prior toperforming the attachment reaction. Where attachment is to take placethrough a linker, the linker can be attached first to the polymer, firstto the biological entity, or concurrently to both. Non-limiting examplesof preferred nucleophile and electrophile groups for use in forming acovalent attachment are presented in Table 1.

Typically, the entities to be covalently attached can be suspended ordissolved in an appropriate solvent, e.g., aqueous methanol, aqueousethanol, acetonitrile, dimethyl formamide, acetone, dimethyl sulfoxide,or a combination thereof, at an appropriate pH, commonly about pH 7 toabout pH 10, and at a temperature from about 10° C. to about 40° C. Aneutral-to-basic pH is typically used and this is in most cases providedby addition of a base to the reaction medium. Examples of preferredbases for this purpose include inorganic bases and organic nitrogenousbases. Among inorganic bases, metal hydroxides, carbonates, andbicarbonates are preferred, preferably alkali metal hydroxides,carbonates, and bicarbonates, and combinations thereof. Examples ofpreferred inorganic bases include sodium carbonate, sodium bicarbonate,sodium hydroxide, potassium carbonate, potassium bicarbonate, lithiumhydroxide, lithium carbonate, potassium hydroxide, and combinationsthereof.

TABLE 1 Exemplary Reactive Group Pairs For Attachment ChemistriesNucleophile Electrophile Attachment Amine Alkyl carbodiimide-activatedester Amide Bromoacetamide Amine Carboxyl Amide Chloroacetamide AmineCyclic carboxylic anhydride Amide 9-Fluorenylmethoxycarbonyl AmideN-Hydroxysuccinimide ester Amide Isocyanate Urea Isothiocyanate ThioureaPhosphate Phosphoramide Phosphonate Phosphonamide Sulfonate SulfonamideAlcohol Alkyl carbodiimide-activated ester Ester (or Thioester) (orThiol) Bromoacetamide Ether (or Thioether) Carboxyl Ester (or Thioester)Chloroacetamide Ether (or Thioether) Cyclic carboxylic anhydride Ester(or Thioester) Ester Ester (or Thioester) 9-FluorenylmethoxycarbonylEster (or Thioester) N-Hydroxysuccinimide ester Ester (or Thioester)Maleimido Ester (or Thioester) Semicarbazido Ester (or Thioester)Thiosemicarbazido Ester (or Thioester) Alkyl tosylate, mesylate, Ether(or Thioether) brosylate, nosylate, nonaflate, triflate, or tresylatesalts

In some embodiments, conducting polymers can be any non-conductivemonomer or polymer that can be made conductive in the presence of anappropriate doping system. In some embodiments, conjugated polymersdescribed herein can also be chemically synthesized to containfunctional side groups that can allow for binding of proteins, lipidsand nucleic acids before or after polymerization. In addition tofunctionalization of the conducting polymers, bioactive molecules,including proteins, lipids and nucleic acids can be also attached to theconductive polymers through hydrogen bonding, electrostatic andnon-polar interactions. In some embodiments, the conductive polymer isbiodegradable and will dissolve in the presence of biological fluid, forexample, when the device is implanted in situ e.g. implantable brainprostheses, neural stimulators, transient heart devices and the like.The biodegradable conducting polymer can include, but is not limited to,polypyrrole poly(3,4-ethylenedioxythiophene) block PEG, andpoly(3,4ethylenedioxythiophene), tetramethacrylate and others which arecommercially available from TDA Research Inc., Wheat Ridge Colo., USA.

Conductive polymers contemplated by the present teachings typicallyrequire counter ions for polymerization and electroconductivity acrossthe electrode-tissue interface. The conducting polymers are reached witha polyelectrolyte at the molecular level. Electron delocalization is aconsequence of the presence of conjugated double bonds in the conductingpolymer backbone. To make the conducting polymers electricallyconductive, it is necessary to introduce mobile carriers into the doublebonds, this is achieved by oxidation or reduction reactions (called“doping”). The concept of doping distinguishes conducting polymers fromall other kinds of polymers. This process can be assigned as pdoping orn-doping in relation to the positive or negative sign of the injectedcharge in the polymer chain by analogy to doping in inorganicsemiconductors. These charges remain delocalized being neutralized bythe incorporation of counter-ions (anions or cations) denominateddopants. In certain embodiments, ionic electrolytes or dopants used topolymerize conducting polymers include but are not limited to:poly(styrene sulfonate) (PSS; Sigma Aldrich, St. Louis, Mo., USA),LiCIO₄, Phosphate-buffered saline (PBS; HyClone, Logan, UT), Hank'sBalanced Salt Solution (HBSS, HyClone), Collagen, Poly-D-Lysine (PDL),Poly-L-Lysine, poly-ornithine, and bioactive molecules of interesthaving the appropriate ionic charge for the type of doping system usedand can include, but is not limited to: dexamethasone or otheranti-inflammatory agents, antibiotics, anti-mitotics, growth factors,scar-reducing, poly acrylic acid, dodecylbenzene sulfonic acid (DBSA),p-toluenesulfonic acid (p-TSA) and combinations thereof.

III. Biological Components

The devices, electrodes and coatings for electrode-based devicescontemplate the use of one or more biological components. The term“biological component” is a term that can encompass any organic materialincluding a complex material such as an agglomeration of cells such astissue to the unitary, such as to cellular constituents, for examplecell structures such as receptors, ion-channels, membranes, organelles,enzymes, antibodies and chromasomes and polymers of amino acids, sugarsand nucleic acids found within, on or produced by cells. In variousembodiments, bioactive molecules can be added to the bioelectrode oradded to the hydrogel scaffold used to support the growth of cells,tissues and other biological components. Bioactive molecules can be anynaturally cell produced protein, lipid, carbohydrate or nucleic acidmolecule that can affect any one of the parameters of expression,differentiation or growth of any biological component, but can alsoinclude natural and synthetic molecules that can affect the sameparameters and can include drugs, pharmaceuticals, biologics andchemicals known to affect such cellular parameters in both prokaryoticand eukaryotic cells. Thus, for the purposes of the present teachings, abiological component can include, but is not limited to: tissue, cellsincluding eukaryotic and prokaryotic cells, archaea, cellularconstituents including membranes of cells, synthetic membranes or filmsmimicking cell membranes with and without membrane proteins includingreceptors, extracellular matrix molecules, e.g., laminin, collagen andfibronectin, receptors, antibodies, ion-channels, proteins,polypeptides, lipids, carbohydrate containing metabolites enzymes, andnucleic acids (RNA, DNA and cDNA) produced by any cell. In someembodiments, an organic living cell can be any living prokaryotic cell,for example bacterial cells, and eukaryotic cells, for example yeast andmammalian cells found in the various tissues, and organs. Organic livingcells can be found in nature or they can be derived from nature andmanipulated recombinantly using methods commonly known in the art tointroduce exogenous DNA and RNA molecules to alter the expression ofproteins and other biomolecules, differentiation and growthcharacteristics of any cell. In some embodiments, the biologicalcomponent can be one or more cells derived from electroactive tissue,including without limitation, cardiac cells, neural cells comprisingneurons, glial cells and cells that are found naturally in muscle.

In some embodiments, the cells, (eukaryotic or prokaryotic cells such asbacteria can be functionalized by adding functional groups such as RGD,IKVAV, YIGSR peptides, and other functional groups which can becovalently attached to the bacterium, cell or cell part, cell membranes,outer membrane proteins (OMPs), cell surface proteins and the like, orthey can be linked to spacers having bifunctional moities that cansimilarly attach to the bacterium, cell or cell part, cell membranes,outer membrane proteins (OMPs), cell surface proteins and the like. Insome embodiments, the eukaryotic cells (including electroactive cells)and bacterial cells can be recombinantly manipulated to express and/orsecrete a variety of cellular constituents and bioactive molecules thatcan be used to enhance the biocompatibility of the bioelectrode,including, growth and differentiation factors, hormones, enzymes, cellsurface antigens (CD antigens), and ion-channels that can attract and/orsupport the cells around the site of implantation.

IV. Biosensors, Diagnostic Devices and Coatings for Devices/Probes andElectrodes

In various embodiments of the present teachings, the present conductingpolymer structures polymerized in the presence of biological componentscan be applied onto bare electrode substrates, to enhance preexistingelectrode based devices and create enhanced microarray electrodecomponents for sensing, recording and stimulating electrical events inphysiological fluid, tissue, cell culture and non-physiologicenvironments such as those encountered in air sampling and watersampling.

A variety of devices and electrode based systems can be functionallyenhanced through the application of the conducting polymer-biologicalcomponent hybrid material described herein including, withoutlimitation, microelectrode-based neural prosthetic devices, cardiacanti-arrhythmia devices (pacemakers), defibrillators, cochlear, retinalprosthetics, deep brain stimulators and electrode based devicescurrently used to detect organic and inorganic substrates, drugs, andbiologics that rely on the detection of current or changes in impedance,resistance or capacitance or surface energy on the biologicalcomponent-conductive polymer interface.

The binding of a molecule such as a cognate ligand, drug, substrate toits receptor, ion-channel or enzyme entrapped in conducting polymer canbe detected and processed using one or more electrodes and processingcomponents. In some embodiments, enzymes can be embedded inelectrochemically polymerized conducting polymer. In variousembodiments, enzymes contemplated for the fabrication of biosensors ofthe present teachings can include any enzyme which participates in aredox reaction when binding to its cognate substrate to any cofactor,including enzyme classes belonging to oxidases, reductases transferases,oxidoreductases, lyases, hydrolases, ligases, and isomerases. Glucoseoxidases can be used in the biosensors of the present teachings tomonitor glucose. Similarly other medically important enzymes can be usedto monitor their cognate substrate in the field of diagnostics labdevices. The bioelectrode can be part of a biosensing device wherein themolecular event of specific binding between the target analyte and thebiological component results in a physical change in electricalresistance which can be measured using cyclic voltammetry and impedancespectroscopy. Automated systems are known in the art that can read andtransduce the electrical signal obtained from the conducting polymer inresponse to surface energy and resistance changes occurring at thebiological component-conducting polymer and analyte interface to thesensing and processing devices attached to the bioeletrode.

V. Instrumentation and Analytic Tools

The electrodes, electrode based devices and coatings used to modifypreexisting electrodes can optionally include controllers, analyzers andother sensing devices and computers that can be used to control theoutput of electrical current, or voltage. These optional components canalso be used to perform, measure and record electrical events, currentflow, electrical impedance spectroscopy, cyclic voltammetry, resistance,conductance, capacitance, and potential of the integrated network to theflow of electrons. These analytical systems and devices are commerciallyavailable, for example the Brinkman's (Eco Chemie) Autolab systemconnected to various CPU's (Windows or Macintosh computers) availablefrom Brinkman Instruments Inc., Westbury, N.Y. USA.

VI. Power Sources, Controllers and Analysis

The bioelectrode, electrode based devices and coatings used to modifypreexisting electrodes can optionally include power sources, actuatorsand controllers for the delivery of current and/or voltage forelectropolymerization of the conducting monomers around the biologicalcomponent and for delivery of current and voltage to the conductingpolymer in proximate contact with the biological component. Powersources can provide voltage potentials, AC and DC current. In someembodiments, the bioelectrode and electrode based devices employing thebiological component-conducting polymer coatings can also be poweredwith batteries.

B. Methods of Manufacture

I. Living Cell Bioelectrodes and Electrode Coatings

Electrodes can be implemented in any aqueous or liquid-saturatedenvironment, such as in living tissues, in the blood stream, in a lake,river, or ocean, in a complex chemical solution, or in most types ofgel. Bioelectrodes can perform a number of activities including 1)direct electrical stimulation or recording of/from a small population ofcells, a single cell or a highly localized region of a cell's membrane;2) extracellular electrical stimulation or recording of cellpopulations; 3) delivery (passive or temporally controlled) of moleculessuch as proteins (growth factors), drugs, chemicals, vitamins, toxins,drug-delivering vesicles/particles; 4) sensing/detection of mechanical,electrochemical and/or biochemical information from the environment.Information gained through the recording or sensing features can berelayed to a computer directly through the bioelectrode apparatus orindirectly via other electrodes connected to the sensing devicesattached to the recorders and computers. In some embodiments, livingcells can be placed in contact with a first electrically conductivesubstrate such as a metal probe or wire. Biological components,including cells can be grown on the first electrically conductivesubstrate, alternatively, the first electrically conductive substratecan be inserted or contacted with living tissue or a solution comprisinga biological component to produce a biologically interfaced electrode.

As shown in FIG. 1, a solution of monomer for example EDOT 30 isimmersed with a biological component which can be any living celladhered onto an electrode (anode as shown in FIG. 1) The conductingpolymer 30 component of the bioelectrode is mixed with an appropriatedopant 34 to provide the necessary redox conditions for polymerization.After an application of galvanostatic or potentiostatic current, themonomers are polymerized with respect to the current emitted from theelectrically conductive substrate as shown in FIG. 1 as the anode.Conducting polymer 50 can be deposited directly onto the surface of theconducting substrate/electrode in such a way that any electrical chargetransmitted to the electrode is transmitted through the conductingpolymer as well, thus the conducting polymer matrix becomes theelectrode itself. In some embodiments, the conducting polymer, forexample, poly(3,4-ethylendioxythiophene) (PEDOT) can be used to preparethe bioelectrode. Bioelectrodes of the present teachings can have lowbiodegradability, low electrical impedance, long-term electricalstability under in aqueous solutions, is mechanically soft, can betailored to have a variety of surface morphologies (varying levels oforder, porosity and roughness at the nanometer and/or micrometer scale),and can incorporate natural or synthetic bioactive molecules/proteins(drugs, chemicals, vitamin, growth factors, cell adhesion proteins,signaling proteins, enzymes, substrate for enzymes) in a spatiallycontrolled manor and if desired these molecules can be released in atemporally controlled manor by the application of low electricalcurrents, and also can incorporate micro or nanoparticles ordrug/molecule delivering vesicles.

The biological components, including living cells that in someembodiments can also include electroactive cells such as heart, brain,CNS and muscle cells can be fully integrated into the conducting polymermatrix in such a way that the plasma membranes of the cells areintimately interfaced as shown in FIG. 2 with the polymer moleculesallowing for seamless electrical signal transduction between theconducting polymer, the cells, and the electrode. The type(s) of cellsincorporated into the bioelectrode device or electrode coatings can betailored according to the desired function of the resultingbioelectrode. For example, in some embodiments, the use of microalgaecan be incorporated as the biological component to conductometricallydetect heavy metals and pesticides in an aqueous environment. In otherembodiments, the biological component can be growth factor-secretingneural stem cells to be incorporated into the bioelectrode or electrodecoatings to promote tissue regeneration and neurite extension towardbioelectrode devices such as a neural prosthetic device implanted in thebrain.

II. Cell Templated Electrodes and Electrode Coatings

In various embodiments highly biomimetic electrodes and electrodecoatings for preexisting electrode substrates and for microarrayelectrode based devices having cell features/cell surface features thatare templated and patterned with nanometer and micrometer scale surfacefeatures are described. The conductive polymers according to the presentteachings are capable of being cast into films of varying thicknesses.Monomers of the conductive polymer can be electrochemically polymerizedon the electrode. In certain embodiments of the present teachings,conductive monomers are polymerized in the presence of living cellsgrown on the surface of the electrode. Cells which can be entrapped inthe conductive polymer can be electroactive cells, (for example, but notlimited to, neurons, skeletal and cardiac cells) capable of conductingelectrical signals or cells capable of interacting ionically with thesurrounding environment. Target cells in the surrounding environment canchemically react with the embedded cells within the conductive polymerto facilitate signal transduction or other chemical redox(reduction/oxidation) reactions for example ligand-receptor andligand-enzyme binding events for example as in the glucoseoxidase/glucose reaction.

In certain embodiments, the conducting polymer network with cell shapedholes/imprints can be created by electrochemically depositing conductingpolymer in the presence of live cells cultured in monolayers on thesurface of probe cell-templated electrically conductive substratefollowed by removal of the cells. The porous or textured cell-basedconductive electrode can encourage cells for example, neurons ormyocytes, to intimately interact with the electrode surface due to thehigh surface area of the fuzzy conductive polymer and the cell shapedimprints. As shown in FIG. 3 panels A-C, the cell-defined polymertopography includes cell shaped holes and imprints as well asmicron-sized tunnels, crevasses, and caves in the polymer surface due tofor example in the presence of neurons, neurite-templated tunnels ofconducting polymer molded around extended neurites that provideinvaginations on the neurite length-scale. In certain embodiments, thecell-templated bioelectrodes can be implanted in tissue, resulting in acell-templated polymer electrode surface that can encourage cells in thehost tissue to re-populate the cell shaped holes and send processes intothe micron-sized crevasses and caves of the polymer surface. Thisintimate contact between cells and conductive polymer at one surface andbetween conductive polymer and an electrode substrate at the othersurface will allow for seamless electrical contact between the electrodeand the target tissue as shown by the electrical signal benefitsprovided by cell-templated bioelectrodes coated .

The bioelectrodes having cell-templated features also increase thecharge transfer capacity of the electrodes and reduce the electricalimpedance experienced by the bioelectrode as compared to a bareelectrode wire, due to the increase in effective surface area of thebioelectrode. As shown in FIG. 4, Panel (A) cell-templated PEDOT had animpedance plot somewhere between electrically conductive substratescoated with PEDOT and electrically conductive substrates coated withSY5Y cells embedded in PEDOT (live cell bioelectrodes).

III. Space-Filling Hydrogel Biolectrodes and Electrode Coatings

In certain embodiments according to the present teachings, thebiocompatible implantable electrode comprises hydrogel seeded with oneor more biological components, for example, living cells through which aconductive polymer network is electrochemically deposited. As shown inFIG. 5, space-filling bioelectrodes can be implanted into living tissue,for example in the brain. The biocompatible hydrogel 120 orthree-dimensionally cross-linked macromolecules of hydrophilic polymerscan be injected through the skull 90 and can serve as both a nutritiveand physically supportive environment for the living cells 110 and as ascaffold for creation of a diffuse conductive polymer network 130 ofmicrometer and nanometer thin fibers. Once injected into tissue, thecells 110 can be any natural or recombinant cell, for example cellsexpressing growth factors such as trophic factor to attract neurites tothe site of implantation. In certain embodiments, the hydrogel materialscan be exceptionally soft, hydrophilic and “tissue-like” thuswell-suited for coating of biomedical devices making possible low levelsof traumatic injury to host tissue during device implantation.Furthermore, the hydrogel can be supplemented with bioactive molecules,for example, drugs (including, but not limited to anti-virals,anti-microbials, anti-oxidants and anti-inflammatory agents) to inhibitadverse immune system reactions and/or biomolecules (cell adhesionproteins, for example: Integrins, neural-cell adhesion molecule, NCAM;Laminins; Fibronectins; Vitronectins; Cadherins and the like) to promotespecific cell-polymer interactions such as synaptogenesis and nerveguidance. The hydrogel matrix can also be seeded with living stem cellswhich can provide for diverse benefits to the host tissue near theimplant site including growth factor secretion to promote local tissueregeneration, recruitment of endogenous stem cells to the device implantsite, and a source of multipotent progenitor cells to replace cells thatwere injured or killed during the device implantation process. Thebiodegradability of the hydrogel can be controlled to slowly resorb,allowing migrating cells (for example neurons) to penetrate theconducting polymer network, eventually leaving the cells in the hosttissue in direct contact with the bioelectrode. Together, the inclusionof additional modulators of inflammation, chemotaxis/adhesion and growthfactors can serve at least three distinct functions: (1) the factors canameliorate and/or mitigate tissue injury and inflammation, (2) increaseregeneration at device implant sites, and (3) facilitate intimatecontact between the bioelectrode and host cells.

In certain embodiments, living cells can be incorporated into an aqueoushydrogel prior to cross-linking of the hydrogel into a 3D scaffold.According to the methods of the present teachings, many suitablenon-toxic hydrogel compositions can be cross-linked in the presence ofliving cells, including, but not limited to hydrogels comprising calcium(cross-linker) alginate and polyvinyl alcohol (PVA), chitosan,self-assembling peptides and functionalized poly(ethyleneglycol)-poly(L-glycolic acid) (PEG-PLGA). The hybrid conductingpolymer-cell-hydrogel compositions can be prepared by embedding anelectrode in the cell-hydrogel complex for example, but not limited to,by using a conductive substrate such as a platinum, silicon, or goldelectrode substrate coated with the cell-seeded hydrogel which can becross-linked around the electrode, or alternatively, a conductivemicrowire can be inserted into a cross-linked 3D hydrogel scaffoldcontaining cells within the living tissue to be treated.

For polymerization, the hydrogel-electrode complex can be submerged inan electrically-connected reservoir containing the desired monomer aswell as ionic dopants or polyelectrolytes in a saline solution such asPBS or HBSS. Galvanostatic current (typically 0.1-100 μA/mm²) can beapplied to the electrode substrate and the solution using an AutoLabPotentiostat/Galvanostat (EcoChemie) for 1 minute-2 hours.Electrochemical oxidation/reduction of the monomer results in theformation of a diffuse conducting polymer network within thehydrogel-cell complex. In some embodiments, microfluidic monomerdelivery devices can be used to deliver conductive monomer to thecell-hydrogel matrix implanted in the tissue.

As shown in FIGS. 6A and 6B, electrochemical impedance spectroscopy(EIS) and cyclic voltammetry (CV) analysis of space-filling electrodesimplanted into living tissue such as the cochlea of a living guinea pighave shown to effectively reduce impedance as shown in Panel (A) ascompared to bare gold with in PBS. The space-filling bioelectrodes ofthe present teachings shows a decrease in impedance over many orders ofmagnitude relating to the development of a high surface area ofconducting polymers in nutritive hydrogel in the guinea-pig cochlea. InPanel (A) and (B) the hydrogel comprised alginate a crosslinking agentand conducting monomer. The conducting monomer was polymerized in situ.The CV shows an increase in charge storage capacity relating to theformation of the conducting polymer network in the porous hydrogel.

In certain embodiments, the hydrogel can be a scaffold or matrix forliving stem cells or progenitor cells which can promote tissueregeneration and wound healing at the bioelectrode insertion site. Inother embodiments, the hydrogel scaffold can be used as a deliverydevice for drugs, proteins and other bioactive molecules, and labelingreagents to the target tissues. In certain embodiments, the bioelectrodecomprising the conducting polymer network within the hydrogel scaffoldcan be used to release drugs or other reagents from within the hydrogelor stimulate differentiation of progenitor cells within the hydrogelmatrix in a controlled manner. In certain embodiments, the conductivepolymer contained within the hydrogel matrix can serve as a mechanicallysoft and non-immunogenic coating. Furthermore, in certain embodiments,the hydrogel can be used as a “space-filling” electrode that would notnecessarily need to be inserted into the target tissue, but rather itcould be placed next to the target yet still allow for electricalinnervation of the target tissue via the growth of the conductivepolymer network contained within the hydrogel.

IV. In Situ Injectable Electrodes

In certain embodiments of the present teachings, the conductive polymernetwork can be directly polymerized within living tissue therebyreducing the likelihood of electrode damage and tissue damage during andafter electrode implantation. In certain embodiments, the resultingconductive polymer network electrode can be in intimate contact with theplasma membrane of living cells. In certain embodiments, the growth ofthe diffuse biological component-conductive polymer hybrid from thesurface of the implanted microfluidic bioelectrode device can create anelectrically-connected diffuse network of molecularly thin polymerfibers and chains woven around cells, effectively innervating thetissue. See FIG. 7. In various embodiments, a vessel containingconducting monomer 150 is immersed with the tissue to be implanted withthe bioelectrode. In various embodiments, the monomer solution 150 canbe injected into tissue 160 for example brain, or heart or musculartissue. To polymerize the conducting monomer 160 in situ, a workingelectrode or first electrically conductive substrate 180 is insertedinto the tissue where the conducting monomer 150 was injected. Next asecond electrically conductive substrate 170 (reference or counterelectrode) is place near the first electrically conductive substrate 180and a constant current is applied to polymerize the conducting polymerin situ. This process can result in the establishment of intimate,specific, and sensitive signal transduction between electrically activecells in the host tissue and the electrode of the implanted deviceresulting in improved electrical charge transfer capacity of theelectrode. See FIG. 7. In certain embodiments of the present teachings,the conducting polymer according to the present teachings can bepolymerized within living tissue resulting in fully integrated andefficacious implanted electrodes for example but not limited to:cortical recording/stimulation, deep brain stimulators, peripheral nerveelectrodes, cardiac anti-arrythmia treatments (bradycardia, tachycardiaand other arythmias), muscle stimulation, surgical ablation (for exampleepilepsy treatments), pH monitoring, glucose sensing, cochlear implants,and retinal prosthetics. In addition, the diffuse, conductive polymerminimizes the necessity of stiff silicon-based or metal based probeelectrodes and signifies a new electrode paradigm built around the supersoft nano-electrode integrated with the living tissue. The diffuseconductive polymer networks polymerized off an implanted electrodedirectly within living tissue can also be independently electricallyconnected to via additional electrodes inserted within the boundaries ofthe conducting polymer network. If desired, the implanted electrode fromwhich the conducting polymer network was originally polymerized can beremoved and a new, independent electrode can be inserted into theconducting polymer network and can then function as the primaryelectrode that interfaces with the conducting polymer network.

V. All polymer Electrodes

In some embodiments, the electrode substrate and all of the implantedcomponents are fabricated with polymeric, non-metallic components. Thepolymer wires/electrodes are non-metallic, non-ceramic, and do notcontain metalloids (e.g. Silicon) or alloys. Polymer electrodes arecomprised of a conducting polymer or combinations of conducting polymersand non-conductive polymers or hydrogels juxtaposed in specificconfigurations resulting in an electrode lead that can be used in placeof “normal” metal electrodes or wires. In some embodiments, the polymerelectrode may also contain carbon or carbon nanotubes. Polymerelectrodes can be used in any situation in which it would beunfavorable, dangerous, or impossible to use metal such as in thepresence of a magnetic field (e.g. MRI scans of individuals withimplanted devices that contain metal electrodes devices orbioprosthetics). Secondly, polymer electrodes can be created severalways from a diversity of substrates and materials and are highlyadaptable and can be readily tailored for specific, diverseapplications.

C. Methods of Use

The electrodes and electrode based device coatings contemplated in thepresent teachings offers the ability to improve electrode performance indiverse electronic biomedical device applications including cardiacpacemakers and defibrillators, biosensors, deep brain stimulators,cochlear implants, retinal prosthetics, and drug detection and bioactivedelivery devices. In some embodiments of the present teachings,multipurpose conducting polymer coatings that are applied to bareelectrodes or preexisting electrode based devices are not onlyelectrically active with low electrical impedance, but they arebiocompatible, mechanically soft, and are “fuzzy” with a high surfacearea at the micro and nano scale thus providing a device surface thatfacilitates direct interactions and seamless integration between theelectrode device and the target tissue or media. Furthermore, thesenovel electroactive conducting polymers can be made bioactive byincorporating living cells, such as stem cells or cells that have beengenetically engineered to express various molecules on their surface, orto produce exogenous bioactive molecules such as growth factors andreceptors, receptors, ion-channels, antigens or antibodies, growthfactors, or other biomolecules as well as can be tailored to have avariety of surface morphologies including nanofibers and nodules,cell-shaped holes, nanosphere-templates, and neurite-templatedmicrotubes.

The novel functions imparted by the electroactive conducting polymercoatings on implanted biomedical devices described herein correspond toreductions in formation and extent of encapsulation of devices infibrous scars, improved ability to record high quality electricalsignals, increased electrical stimulating capacity, and enhanced devicelongevity. In some embodiments, the electrode substrates are coated withthe electroactive biomaterial of the present teachings comprising anelectrically conductive polymer, and a biological component. At leastsome of the electrically conductive polymer is disposed and polymerizedin proximate contact with the biological component and the electricallyconductive substrate.

I. Sensing and Recording Electrodes

Chronic implantation of existing microelectrode-based neural prostheticdevices is associated with CNS injury and inflammation which results inneuronal loss around electrode sites and formation of a high impedanceglial/immune cell encapsulation of the prosthetic device. Together,these phenomena serve to diminish the quality of recordable neuralsignals over time following implantation, eventually blocking thecapacity to record. This undermines the ability to establish thebrain-computer interface that is necessary for function of the corticalprosthetic.

The novel bioactive, biocompatible, low impedance electroactiveconducting polymer coatings comprising biological components are anideal surface modification of recording electrodes on neural prostheticdevices. The “fuzzy”, large effective surface area makes for lowelectrical impedance at the electrode-tissue interface which increasesthe probability of recording high quality signals from target cells evenin the presence of a fibrous glial encapsulation.

In various embodiments, sensing and recording electrodes and microarrayelectrodes for neuronal activity mapping require fully integratedbioelectrode devices that can interface with the surrounding tissueintimately. Improved electronic responses of these types of electrodescan be achieved by increasing the effective surface area of theelectrode. The conductive polymers of the present teachings provide suchincrease in effective surface area and can support the growth andsurvival of living cells, for example stem cells expressing growth anddifferentiation factors in the implanted site. The sensing electrodescan further increase selectivity and sensitivity by mimicking naturalscaffolds comprising cells and bioactive molecules including drugs,growth factors, anti-inflammatory agents and antibiotics. In someembodiments, bioelectrodes comprising conductive polymer, polymerizedaround living cells can be implanted into tissue to enable accurate andhigher probability of recording higher quality signals due to theincrease in charge capacity and decrease in impedance of the due toimmune cell encapsulation and glial scar formation around thebioelectrode. Together, this will allow for more stable and sensitivelong-term neural recording than current neural prosthetic electrodetechnology. Electrodes coated with a cell embedded conducting polymernetwork of the present teachings can include cortical prosthetics,advanced catheters for electrophysiological mapping of electroactivetissue such as the heart, CNS, brain and muscle.

In some embodiments, electrodes of the present teachings can also bemodified to provide highly tissue specific and biocompatible coatingsfor the attachment of electrically active tissue. Followingpolymerization of the conducting polymers around one or more biologicalcomponents and the electrode substrate, the three dimensional surface ofthe electrode can be made even more attractive to the cells of thesurrounding tissue by removing the embedded biological component, forexample cells (neurons, myocytes, fibroblasts, stem cells etc) leavingbehind cell membrane components and pores or invaginations andthree-dimensional structure that facilitates the binding andcolonization of adjacent or neighboring cells to the electrode. Theresulting electrode is highly integrated and biomimetic, enablingsensitive recordation of functional characteristics of electroactivecells. The inventors of the present teachings have found significantincreases in charge transfer capacity and marked lowering of electricalimpedance when the electrodes of the present teachings are coated with“fuzzy” cell templated conducting polymer structures. Attracting cellsto the electrode and encouraging the neighboring cells to settle andoccupy the cell shaped holes, tunnels and crevasses left behind afterremoval of the biological component from the electrode substrate servesto improve electrode stability, prevent erosion of the electrodesurface, diminish electrode biofouling due to adverse immune reactionsand improves the performance of the electrode in comparison with hardmetal electrodes currently in use.

In some embodiments, cell recruitment and improved communication betweenthe electrode substrate and the surrounding tissue requires more than acell-templated structure with nanoscale features. For these types orapplications, such as implantable electrodes into the brain, heart, andcentral and peripheral nervous systems, electrode sensing and recordingrequires an even greater degree of biocompatibility and molecularmimicry. In some embodiments, the electrodes of the present teachingoptionally include a hydrogel material that can be implanted into asubject either before insertion of the electrode substrate or can beimplanted concomitantly with the electrode substrate in to the subject.

Hydrogel scaffolds comprising alginate, poly-vinyl alcohol and otherbiocompatible materials can be implanted or injected into the electrodesite prior to insertion of an electrode. In some embodiments, thehydrogel scaffold can be biodegradable or non-degradable. For examplesof hydrogel scaffolds for use with conducting polymers see Gilmore, K.et al., Polymer Gels and Networks, 2: (1994) 135-143, and Ghosh, S. etal., J. The Electrochem. Soc. 147:1872-1877 (2000). The presentteachings provides markedly improved hydrogel scaffolds when in use insitu due to the polymerization of the conductive polymer in the hydrogelin the presence of one or more biological components. In someembodiments, the hydrogel containing the conducting monomer is injectedinto a site for example a cavity, in interstitial spaces and generallyaround cells of interest or within tissue then a first electricallyconductive substrate is inserted from which a conducting polymer networkis polymerized in situ. The conducting polymer network forms around themacromolecules and fibrils that comprise the hydrogel and use thesehydrogel components and features as a scaffold for polymerization in away similar to how the conducting polymer networks form when polymerizeddirectly within tissue. The resultant bioelectrode comprises conductivepolymer embedded around cells within a hydrogel framework. In someembodiments, the hydrogel is supplemented with other cells for example,recombinant stem cells producing neurotrophic growth factors, and otherbiomolecules of interest that can support the growth and development ofthe surrounding cells and tissue.

In some embodiments, recording devices comprising a biologicallyintegrated bioelectrode device can be used to record or detectelectrical signals between cells and between tissues. A method ofelectrically detecting a transfer of electrical signals between livingcells, comprises the steps: providing a bioelectrode device comprising afirst electrically conductive substrate in intimate contact with tissuecapable of transferring electronic charge. The bioelectrode deviceincludes a first electrically conductive substrate; a biologicalcomponent; and a conductive polymer electrically coupling the firstelectrically conductive substrate to the biological component tocollectively define a bioelectrode. The bioelectrode transmits orreceives an electrical signal between the first electrically conductivesubstrate any one of the biological component and the conductivepolymer. The circuit is achieved by electrically connecting thebioelectrode device and a second electrically conductive substrateelectrically coupled with the bioelectrode to a power source. Onceconnected, the power source is applied providing an effective amount ofvoltage or current across the first and second electrically conductivesubstrates, thereby inducing a voltage or current across the conductivepolymer. The system detects the transfer of electrical signals with saidbioelectrode device.

In some embodiments, the bioelectrode device can also include otherelectrodes such as a reference electrode, counter electrode and asaturated calomel electrode during sensing, recording, and stimulatingcells and for polymerizing conducting monomer.

II. Stimulating Electrodes

Conductive polymer coatings containing one or more biological componentssuch as cells, receptors, cell membranes, cell matrix proteins and thelike on electrodes will improve electrical stimulation of cells incontact or in the vicinity of the bioelectrode, including, neurons,myocytes and muscle cells by increasing the charge capacity ofelectrodes, and by extending conductive surface from a planar electrodetowards neurons with fuzzy polymer tendrils. In some embodiments, thebioelectrode and hybrid biological component-conducting polymerelectrode coatings of the present disclosure, can also immobilize drugsor living cells to secrete agents for scar prevention/reduction, improveneuronal viability, attract neuronal processes and to promoteintegration between tissue and the bioelectrode for stable and directsignal transduction. In some embodiments, a bioelectrode comprising afirst electrically conductive substrate seeded with neuronal progenitorcells and coated with electrochemically polymerized conducting polymercan be implanted into a neuron rich tissue, such as the brain or centralor peripheral nervous system to electrically and biologically stimulatethe growth of endogenous neurons. The progenitor neuronal cells cansecrete factors that can attract the resident population of neurons tointegrate and communicate with the bioelectrode. After integration ofthe neurons as evidenced by morphological extension of micropodia andneurite outgrowth towards the bioelectrode, the bioelectrode can beactuated by applying a voltage and/or current bias to stimulate thegrowth of interconnected neurons in communication with the bioelectrode.Without being limited to any particular theory, improvedbiocompatibility between the bioelectrode and the surrounding tissue canalso be attributed to the reduction of inflammation and glial scarformation, facilitation of tissue regeneration near the implanteddevice, and formation of intimate contact between the electrode surfaceand target neurons.

In some embodiments, the bioelectrodes of the present teachings cansimilarly affect the electrical function and well being of otherelectrically active cells, including cardiac cells or electricallyresponsive cells such as fibroblasts, and bone forming cells. Hybridconducting polymer-biological component coatings polymerized onelectrode substrates can provide direct integration betweenelectrically-active cardiac muscle cells and implanted electrodes.Electrode coatings comprising fuzzy, soft fibril conducting polymerscontaining biological components, for example cells or cell membranes,ion-channels, receptors growth factors and enzymes provide increasedstimulating charge capacity, while decreasing recording electrodeimpedance. In some embodiments, cells that are responsive to electricalstimulation such as those involved in wound healing and for therapiesrelated to bone healing and growth can be particularly benefited bystimulation therapies provided by implantable bioelectrodes of thepresent teachings as a therapeutic form of treatment. Living cells canbe grown or otherwise immobilized in natural or synthetic hydrogelscaffolds and fitted with an electrode substrate, for example, a wire orprobe before implantation to improve and direct the cellular response tothe bioelectrode device and facilitate integration with active tissue.In some embodiments, the living cells embedded in the conducting polymercan include any therapeutically beneficial cell, including cellsrecombinantly made to express and secrete growth factors,differentiation factors (including one or more of Insulin-Like GrowthFactor, NGF family of neurotrophic factors, ciliary neurotrophic factor(CNTF), and pituitary adenylate cyclase-activating peptide (PACAP), BoneMorphogenetic Proteins 1-17, Fibroblast Growth Factor, and any commonlyknown growth factors used to regenerate neural, cardiac, bone and muscletissue) receptor agonists and/or antagonists, enzyme inhibitors andother therapeutically effective bioactive agents known to beadministered to subjects having diseases and conditions of theelectro-active tissue, including the heart, the brain, the central andperipheral nervous system and muscular system.

In various embodiments, tissue regeneration and can be facilitated byimplanting bioelectrodes comprising conducting polymers polymerizedaround embryonic and/or hematopoetic and/or parenchymal stem cells thatare capable of differentiating into neurons, muscle cells and cardiacmyocytes in a hydrogel scaffold containing anti-inflammatory agents andother growth and differentiation factors. In some embodiments, thebioelectrode can release stored drugs and other bioactive substances,into the tissue-electrode interface, particularly when used ascounterions or when added to hydrogel scaffolds. In some embodiments,the nutritive hydrogel scaffold surrounding the bioelectrode can protectthe embedded biological component in the conducting polymer and/orhydrogel on or around the bioelectrode from the immune system andprovide for their growth and differentiation. Since the hydrogelscaffold containing polymerized conducting polymer is in intimatecontact with the electrode substrate through polymerized conductingpolymers, electrical therapy may be administered to program the immatureelectro-active cells for adult cell function.

III. Method of Marking Location of Electrode

In some embodiments, the placement of electrodes including micrometerthin wires and other electrode substrates can make subsequentlocalization difficult, particularly, if the electrode is devoid of anymetallic material or has been removed prior to histological analysis ofthe tissue. In some embodiments, the bioelectrode, electrode coatings,and in situ polymerized biological component-conducting polymer materialdescribed herein can be to modify, preexisting electrodes, can provide avisual cue as to the location or in some cases the former location ofthe implanted electrode in the tissue. The conducting polymer istypically well contrasted when implanted into living tissue.

IV. Method of Attaching an Electrode to Implanted Tissue

In some embodiments, preexisting electrodes and new electrode devicescan be made secure or anchored within the implanted tissue by in situpolymerizing conductive polymer networks within the surrounding tissuefrom the implanted electrode. This method provides for a “fuzzy”three-dimensional architecture that sends nanoscale fiber, fibrils andother structures into the interstitial spaces of the tissue, thusanchoring the electrode in the implanted tissue. This alleviates theproblem of electrode slippage and movement during the period of time theelectrode is implanted and can secure the location of the electrode toits proper stimulating or recording site.

V. Organism, Cell, Drug, Chemical, and Biomolecular Sensors

Electroactive conducting polymers can be used to immobilize detectingagents including but not limited to live cells, cell components, nucleicacids, molecule-functionalized micro-nano particles, enzymes, proteins,or peptides on the surface of electrodes without compromising theinherent properties of the conducting polymer which can act as sensingdevice due to electronic reactions that occur within the polymer whenelectrical current is applied. This behavior can be exploited to detectoxidation/reduction reactions taking place near the surface of thepolymer or within the conducting polymer matrix. Paired with the abilityto incorporate a diversity of detecting agents into the conductingpolymer matrix, this provides a powerful system for sensing electrontransfer reactions that occur between a detecting agent and itscomplimentary counterpart for example between enzymes and theirsubstrates, receptors and ligands, antibodies and antigens, or cells andpathogens. By sensing the electronic transfer of the immobilizedmolecules, electroactive conducting polymer electrodes have been shownto detect concentration fluctuations of many molecules includingglucose, choline, phosphate ions, nucleic acids, and chlorpromazine anddopamine.

In addition, the conductive polymer can be readily modified to contain avariety of bioactive agents to facilitate interactions with specificproteins or biomolecules and limit non-specific interactions that areassociated with device surface biofouling. Proteins can be incorporatedinto conducting polymer films via a variety of methods such aselectrochemical deposition, covalent linkage, and entrapment in theconducting polymer network. This feature of conducting polymers can beexploited to make the conducting polymer network embedded with one ormore biological components, bioactive as well as to make possiblereversible changes in electrical conductivity triggered by specificstimuli thus allowing the bioelectrode to act as a biomolecule sensingdevice.

In some embodiments, electrodes comprising conducting polymerspolymerized around cells or specific biological components such asreceptors, antibodies, ion-channels can be used to detect specificchemical entities. In some embodiments, biological cells, embedded withconducting polymer present on electrode substrates can use the nutrientsin the environment to maintain viability while performing a biosensingfunction by constantly sampling the environment to detect specific(pre-determined by cell type used) biochemical or electrochemicalchanges. If changes in the environment are detected, the signal istransduced by cell-associated enzymatic reactions which result in localalterations in the net surface charge on the cell or receptor which isthen transduced to the electrode via the conducting polymer on thehybrid conducting polymer-biological electrode.

In some embodiments a biosensor can be used to detect a biologicalmaterial in fluid. The method used to manufacture such a device caninclude combining and placing a first and second electroconductivesubstrate on a support. The support can be any biocompatible materialthat is not subject to degradation such as biocompatible plastics e.g.Teflon, ceramics, e.g. porcelain and metallic materials, e.g. stainlesssteel. A solution of biological component for example, an enzyme or cellreceptor or antibody is applied to a portion of the firstelectroconductive substrate. The biological component can have aprotective, porous and fluid transmissible agent including abiocompatible hydrogel, for example alginate hydrogel to form a layer onthe first electrically conductive substrate. Conducting monomer is addedto the solution comprising biological component and hydrogel, and arehomogenously mixed over a portion of the first electrode substrate. Theconductive monomer is then polymerized either by elelctrochemicalpolymerization by applying a galvanostatic or potentiostatic current tothe first electrically conductive substrate or by direct oxidativepolymerization, to form a network comprising conducting polymer aroundthe biological component in a hydrogel matrix. A receptacle is prepared(which can include any vessel capable of holding a solution and twoelectrically conductive substrates, for example a cuboidal flow cellmade of non-conductive material), containing a sample comprising atarget analyte to be analyzed. In some embodiments, the target analytecan be any substance whose presence or quantity is to be determined. Thetarget analyte is a binding partner to the biological component, forexample a specific antigen to its antibody, a specific ligand to itsreceptor. The first and second electrically conductive substrates areplaced in the receptacle. The applied potential is selected to drive anelectronic charge transfer including electron transfer between thebiological component and the conductive polymer.

Specific binding of the target analyte to biological component resultsin a measurable potentiometric or amperometric electronic chargedifference on the surface of said biological component which istransduced to the conducting polymer which is in intimate contact withthe first electrically conductive substrate. The current generated as aresult of electronic charge transferred from the biological component tothe conductive polymer, then to the first electrically conductivesubstrate will be directly proportional to the concentration of thetarget analyte thus allowing for quantification of the concentration ofthe target analyte. A biasing source applies a constant potentialbetween the first and second electrically conductive substrates when thedevice is in the receptacle in contact with the fluid containing thetarget analyte.

In some embodiments, the bioelectrode can contain a thinnon-biodegradable hydrogel coating around the hybrid cell-conductingpolymer matrix of the bioelectrode to prevent exposure of the biologicalcomponent with the external environment. Similarly, in still furtherembodiments, the hybrid cell-conductive polymer electrode can bemaintained by providing a source of nutrition to the embedded cells onthe electrode. In order to protect the nutritive gel from a potentiallytoxic environment, an additional thin layer of a non-resorbable hydrogelcan be used to protect the nutritive gel from degradation. Providing asource of nutrition to the embedded cells contained and coated withconducting polymer makes it possible for the cells of the bioelectrodeto interact (e.g. detect biochemicals or secrete drugs) with theenvironment without actually being exposed to the environment andwithout exposing cells/tissues in the environment to the cells of thebioelectrode of the present teachings

A novel conducting polymer sensor coating for an electrode-based drugdelivery device allowing integration of real-time sensing of the targetmolecule with feedback to the drug delivery device then stimulation ofcontrolled drug release. A bioactive molecule sensing/monitoring orchemical sampling device that makes possible quantification of even asingle target bioactive molecule in a solution. A “smart” polymersurface that can detect when it is in contact with a specific cell typedue to an enzyme-mediated sensing reaction that occurs when a ligandwithin the polymer binds its target receptor. This process could be usedto deliver cell-type specific signals from the polymer film.

VI. Bioactive Catheters

To prevent clogging/clotting around implanted catheters, current “smart”catheters use drug-eluting polymer coatings to prevent cell and proteinadhesion and encapsulation. The conductive polymer-biological componenthybrid polymer coatings of the present teachings can reduce cellular andnon-specific protein adsorption through controllable surface charge,controlled reversible shape-transformations, incorporation of bound orreleasable drugs or proteins, and by immobilization of living cellswhich can release therapeutic agents to direct integration of the devicewith the surrounding tissue.

In some embodiments, the electrode coatings and associated bioelectrodedevices that are contemplated by the present embodiments have resultedin the ability to interface conducting polymer with a biologicalcomponent, for example, plasma membrane of living cells. The presentbioelectrode devices can be incorporated as a new type of material forembedding and encasing living cells to facilitate studies on cellsurface features due to the ability to form a 3D “negative” image of thecell and on real-time dynamics of the plasma membrane polarization andion channel activity throughout the cell regions (soma,dendrites/processes). A novel material for immobilizing living cells ona substrate for no or low vacuum microscopic imaging (TEM, AFM, ESEM,EFM) and possibly for other surface analysis techniques that have yet tobe used on live cells such as FTIR and SFG.

In some embodiments, the conducting polymer-based microelectrode array(MEA) devices of the present teachings comprising one or more biologicalcomponents operable to receive an electrical signal and an electricallyconductive substrate for receiving the electrical signal; and aconductive polymer matrix comprising a plurality of conductive polymersdisposed and polymerized adjacent to the electrode and an biologiccomponent, wherein at least some of the plurality of conductive polymerstransmitting the electrical signal between the electrode and thebiological component can be used for electrical stimulation andrecording of cellular action potentials and extracellular fieldpotentials from single or multiple living electrically active cells. Theintimate contact between the conducting polymer and the plasma membraneof cells allows for sensitive, highly localized, even sub-cellularstudies on synaptic communication and activation of neural activitywhich is not possible with currently available MEAs or common patchclamp-based electrophysiologic techniques.

In various embodiments, the bioelectrode can be used in a method forvisualization and analysis of cell-substrate adhesions made possible bythe inability of the conducting polymer to form on areas of thesubstrate on which cell membrane is adhered. This reveals the details ofthe cell-substrate adhesions at the nanometer scale. Use ofelectroactive biomaterials for these studies would be a cheap, quickalternative to current methods including immunocyto/histochemistry andtotal internal reflection microscopy (TIRF) which are widely used by thebiomedical research community. Quartz crystal microbalance analysis canalso be used to assay cell-substrate adhesion but this method is notwidely available to biomedical researchers.

The electrode compositions of the present teachings can replace similarbiomedical devices implanted in other peripheral tissues of the body.

In certain embodiments of the present teachings, cell-based conductivepolymer electrodes are soft, fuzzy materials with low electricalimpedance and enormous effective surface areas. The large effectivesurface area of the conductive polymers can facilitate maximal chargetransfer between electrode and target environment. Furthermore, incertain embodiments, the pliability of the conductive polymer can allowfor decreased mechanical strain at the interface between the soft tissueand the hard device surface compared to a metal electrode substratealone.

Furthermore, the bioelectrode can be inserted and implanted in theinterstitial spaces in the tissue and in the extracellular matrixbetween cells resulting in an electrode that can be intimatelyintegrated with cell surfaces yet due to its molecular and nanometerscale, should not trigger an immune response.

The present disclosure will be further understood with reference to thefollowing non-limiting examples.

EXAMPLES Example 1 Bioelectrodes Comprising Living Cells

Preparation of electrodes and cell cultureware for cell culture: Cellsare adhered to or cultured on conductive substrates or electrodes forthe electrochemical polymerization process. The electrode is sterilizedprior to exposure to cells by washing in 70% ethanol (Sigma-Aldrich, St.Louis, Mo.) for 10 minutes or exposure to UV light for 20 minutes. Thesterile electrode is then placed in a dish for culturing the desiredtype of cells. Depending on the geometry of the electrode in some casesit is necessary to affix the electrode to the bottom of the cell culturedish to reduce lateral movement of the electrode. This is accomplishedby gluing the electrode to the bottom of the dish using a minimal volume(e.g. 1-10 ul) of superglue that is adherent in a liquid environmentsuch as Loctite Cyanoacrylate (Henkel Corp., Rocky Hill, Conn.). Formost cell types, to allow for cell adhesion, the height of the electrodemust be no more than 250-500 um above the surface of the cell culturedish and the electrode surface area should be large enough that a numberof cells can be seeded on the conductive surface (e.g. >100 um²). Thecell culture dish must have a charged surface to promote cell adhesion,thus we use plasma-treated cultureware from Corning (Corning, N.Y.) andpoly(lysine)-coated cultureware (BD Biosciences, San Jose, Calif.). Inaddition, for some cell types (such as neurons) it is necessary to alsocoat the electrodes in poly(lysine) to allow for cell adhesion to thesurface. To do so, the electrodes are first sterilized then glued to thebottom of plasma-treated cell cultureware then a solution ofpoly(lysine) (1 mg/ml; Sigma-Aldrich) in phosphate buffered saline (PBS;Hyclone Media, Kansas City, Mo.) is added to the dish and allowed toincubate for 2-12 h at room temperature (RT) under sterile conditions(in the tissue culture hood). After incubation, the poly(lysine)solution is rinsed with one wash of PBS, then cells can be plated.

Tissue/cell culture: The cell type of choice is maintained in cultureaccording to published methods appropriate for that cell type. Forexample, in current studies our laboratory is using the SH-SY5Yneuroblastoma-derived cell line (gift of Dr. Eva Feldman at theUniversity of Michigan; also available at the American Tissue CultureCollection, www.atcc.org) as well as dissociated cortical neuronalcultures from embryonic mice. SY5Y cells are maintained in Dulbecco'sModified-Eagle's Media (DMEM with glucose, with L-glutamine;Gibco/lnvitrogen, Carlsbad, Calif.) supplemented with penn-strep mixedantibiotic solution (dilute 1:100 in cell media; Gibco/lnvitrogen) and10% fetal bovine serum (FBS; Gibco/lnvitrogen). The media is changedonce per week and cells are passaged/split 1:4 every 2 weeks.

Electrochemical polymerization in the presence of the living cells: Togenerate the living cell bioelectrode, the cell-seeded electrodesubstrate is placed in an electrically-connected reservoir containing anaqueous solution (depending on cell type) such as water, PBS or HBSSthat contains the desired monomer with ionic dopants and/orbiomolecules. Galvanostatic current is applied to the electrode and themonomer solution using an AutoLab Potentiostat/Galvanostat (EcoChemie,The Netherlands) or some similar instrument capable of delivering directcurrent (DC) at 1-10 μA/mm² for 0.5-10 minutes depending on the desiredthickness of the conducting polymer film. Electrochemicaloxidation/reduction of the monomer results in the formation ofconducting polymer films and networks around and onto the adhered cells,thus embedding and immobilizing them in a conductive polymer scaffold.To generate the living cell bioelectrode, the electrochemicalpolymerization can be optimized for each cell type and electrodeconfiguration so that the resulting hybrid cell-conducting polymermaterial is such that the polymer surrounds the living cells and theirprocesses but does not cover the entire cell body.

Cell maintenance in the hybrid cell-polymer electrode matrix: In orderfor the living cell bioelectrode to function properly, the cellsincorporated into the hybrid cell-conducting polymer matrix shouldremain viable for the length of time that the device is expected tofunction. Living cells require access to a host of nutrients, growthfactors, and dissolved gases that are specific to each cell type.

Characterization of surface morphology: The surface morphology of theliving cell bioelectrode can not be characterized without destroying theintegrity of the “living” electrode 2 batches of electrodes wereprepared, 1 batch for electrical characterization and experimentationand another batch for microscopic evaluation. The bioelectrodes areevaluated microscopically to assess cell viability, cell morphology, andintegrity/quality of the hybrid cell-conducting polymer matrix usingoptical and fluorescence microscopy. In addition the surfacetopography/features are explored using AFM in tapping mode in an aqueousenvironment and as well as environmental scanning electron microscopy(ESEM) which is performed in a very low vacuum on a chilled stage(Peltier stage) with 50-70% humidity in the chamber.

Optical microscopy is conducted with a Nikon Optiphot POL, having thecapability for both reflected and transmitted light observations. Imagesare acquired with a Spot RT digital camera running on a Macintosh G4computer. For fluorescent microscopy we use Olympus IMT-2 upright lightmicroscope with Hoffman modulation contrast and a Leica DMIRBfluorescent inverted microscope both with mercury arc lamps for UVlight, Olympus CCD cameras, and accompanying Olympus digital imagingsoftware running on Dell PC computers. Information about the samplesurface topography will be obtained via AFM with a Digital InstrumentsNanoscope IlIl with a Multimode head, located in the Michigan ElectronMicrobeam Analysis Laboratory (EMAL). The images obtained consist of512×512 arrays of height data over scan sizes typically ranging from 100microns down to 1 micron. Information about the surface and themicrostructure of the living cell bioelectrodes can be obtained usingthe FEI Quanta 200 3D Focused Ion Beam Workstation and EnvironmentalScanning Electron Microscope.

Assessment of cell viability: In order for the living cell bioelectrodeto function properly, the cells incorporated into the hybridcell-conducting polymer matrix should maintain viability once embeddedin the polymer as well as throughout the lifetime of the device. Cellviability can be assessed using a variety of methods, many of which arecell type specific. Two assays can be used that are common to many typesof mammalian cells; the Vybrant Live/Dead Assay (Molecular Probes,Eugene OR) and immunocytochemistry for cell death associated proteins,specifically the apoptosis-associated protease activated caspase 3(antibody available from Cell Signaling Technologies, Beverly, Mass.).For the Vybrant Live/Dead assay cell quantity, size, and type of nuclearstaining intensity are measured using 3 different dyes, specificallyHoechst (permeable to all cells but brighter in nuclei of dying cells),SYTOX green (mostly present in cells dying by apoptosis) and propidiumiodide (present in any cell with a compromised membrane; apoptotic ornecrotic cells). For immunocytochemistry (ICC), cells are fixed in 3.7%formaldehyde diluted in PBS for 30 min at room temperature (RT) orovernight (ON) at 4 C. Then cells are washed in cold PBS, thenpermeabilized for 1 h to ON in PBS+0.1% Triton-x (PBSX). Non-specificlabeling is blocked by incubating cells for 1 h at RT in 3-5% bovineserum albumin (BSA)+PBSX. Cells are then exposed to the primary antibody(in this case anti-activated caspase 3 @ 1:100) for 2 h-ON diluted in1.5-3% BSA/PBSX or BSA/PBS. Next cells are washed 3 times in PBS orPBSX, then incubated with the fluorophore-conjugated secondary antibody(1:100 in 1% BSA/PBSX) in the dark at RT. Cells are then washed 3 timesin PBSX, all nuclei are counterstained with Hoechst 33342, then cellsare mounted in Vectashield aqueous mount and stored at 4° C. untilmicroscopic imaging.

Cultured cells can include, but are not limited to: fibroblasts,neurons, myocytes, smooth muscle, glia, Schwann cells, progenitor cells,embryonic stem cells, neural or other stem cells can be cultured onelectrode substrates for the electrochemical polymerization process. Theelectrode can be sterilized prior to exposure to cells by washing in 70%ethanol (Sigma-Aldrich, St. Louis, Mo.) for 10 minutes or exposure to UVlight for 20 minutes. The sterile electrode can be fixed to the bottomof a cell culture dish using a minimal volume (e.g. 1-10 ul) ofsuperglue that is adherent in a liquid environment such as LoctiteCyanoacrylate (Henkel Corp., Rocky Hill, Conn.). To allow for celladhesion to the electrode substrate, the height of the electrode can beno more than about 250-500 um above the surface of the cell culturedish. The electrode substrate surface area should be large enough that anumber of cells can be in contact with the conductive surface (e.g. >20um²). The cell culture dish surface can be charged to promote celladhesion, (plasma-treated culture ware from Corning (Corning, N.Y.)) andpoly(lysine)-coated cultureware (BD Biosciences, San Jose, Calif.). Inaddition neurons can be cultured on electrode substrates, after coatingthe electrode substrate with poly(lysine) to allow for cell adhesion. Todo so, the electrodes can be sterilized first then glued to the bottomof plasma-treated cell cultureware then a solution of poly(lysine) (1mg/ml; Sigma-Aldrich) in phosphate buffered saline (PBS; Hyclone Media,Kansas City, Mo.) is added to the dish and allowed to incubate for 2-12h at room temperature (RT) under sterile conditions (in the tissueculture hood). After incubation, the poly(lysine) solution is rinsedwith one wash of PBS, then cells can be plated.

Tissue/cell culture: The cell type of choice can be maintained inculture according to published methods appropriate for that cell typeand as described above. Tissue culture methodologies and materials canbe found in Coligan, et al., Current Protocols in Immunology, Wileylnterscience, 1991 and is hereby incorporated by reference.

Primary neuronal cultures can be prepared from timed pregnant mice(Swiss-Webster) which can be ordered to arrive on embryonic day 13-14.On embryonic day 18,19, or 20, the mouse is sacrificed by CO₂asphixiation, the embryos are removed and immediately placed in ice coldHank's buffered saline solution (HBSS without MgCI or CaCI; HycloneMedia). The embryos are decapitated and the brain is removed from theskull, the meninges are removed, and the neocortex is dissected. Thecortical tissue pieces are placed in 45 ml of ice cold HBSS untildissociation (usually no longer than 20 minutes). The tissue pieces arewashed 3 times in fresh ice cold HBSS using a sterile 10 ml pipette totransfer the tissues. Cortical tissues from 10-15 mouse embryos aresubmerged in 2 ml dissociation media. The dissociation media is composedof Neurobasal media (Gibco/lnvitrogen) supplemented with 0.5 mML-glutamine (Gibco/lnvitrogen), 5% FBS, and penn-strep (dilute 1:100 inmedia). A 1000 mL pipette tip is used to mechanically disrupt/dissociatethe tissues in the dissociation media (dial pipetter to 500 μl andtriturate 20 times, do not generate bubbles). Once the tissues arecompletely dissociated, the cell suspension is centrifuged at 1000×g for3 minutes at room temperature (RT). The supernatant is removed and thecell pellet can be resuspended in plating media. Plating media isNeurobasal media supplemented with 0.5 mM glutamine, penn-strep, 1% FBS,2% B27 serum-free media supplement (Gibco/lnvitrogen). Cells can beplated on poly(lysine) coated cultureware. Every 5-7 days after plating,a 30% media exchange can be performed. Cultures can be ready forexperimental use by 7-10 days and can remain useful for as long as 21days in culture.

Example 2 Cell-Templated Electrodes and Electrode Coatings

Removal of cells to generate cell-templated conducting polymer films:Following electrochemical polymerization on the surface of the electrodesubstrate as previously described in Example 1, the cells embedded inconducting polymer are removed by mechanical disruption by vigorousshaking in water, saline solutions or exposure to calcium chelatingagents (EDTA, EGTA; Sigma-Aldrich) and proteolytic enzymes such astrypsin (Hyclone Media) which cleave proteins that adhere the cells tothe electrode substrate. In certain embodiments, the use of water washesand mechanical disruption in combination, will remove the cell bodiesand most of the cell material but leaves behind some cell membranecomponents and cell adhesion proteins normally present on the cellsurface. The resulting conducting polymer surface has cell-templatedfeatures lined with cell surface and cell-substrate adhesion proteinsand/or protein fragments that can facilitate binding of cells andtissues that contact this bioactive and biomimetic material. Incontrast, cell removal by exposure to proteolytic enzymes results in aconducting polymer film with cell-templated features but with noinherent biological activity due to the removal of all cell material.

Experimental

Materials and Methods: SH-SY5Y neuroblastoma-derived cells weremaintained in Dulbecco's Modified-Eagle's Media (DMEM with glucose, withL-glutamine; Gibco/lnvitrogen, Carlsbad, CA) supplemented withPenn-Strep mixed antibiotic solution (dilute 1:100 in cell media;Gibco/lnvitrogen) and 10% fetal bovine serum (FBS; Gibco/lnvitrogen).Mouse primary dissociated cortical cultures (MCC) were prepared fromembryonic day 18-20 (E18-20) mice. The brains were removed and submergedin ice-cold Hanks buffered saline (HBSS; without calcium chloride,magnesium chloride, magnesium sulfate, or phenol red; Invitrogen), theneocortex was dissected, the meninges were removed, tissue was washed inice-cold HBSS then manually dissociated with a 1 ml pipette tip. MCCwere maintained in Neurobasal media supplemented with 0.5 mM L-glutamineand 2% serum-free nutritional supplement B27 (Invitrogen) at 37° C. in5% CO₂. A third of the media was replaced every 4 days, and cells wereallowed to mature for at least 7 days before use in experiments.

Electrodes for cell culture: Prior to exposure to cells, electrodes(bare or PEDOT-coated) were sterilized by washing in 70% ethanol(Sigma-Aldrich, St. Louis, Mo.) for 10 minutes. We used two differenttypes of electrodes for these studies, custom-designed, in-housefabricated Au/Pd sputter-coated electrodes (Au/Pd; 6 mm diameter) andApplied BioPhysics (Troy, N.Y.) ECIS electrodes (ABP; 250 μm diameter).For the Au/Pd electrodes, it was necessary to glue the electrode to thebottom of the cell culture dish to prevent lateral movement of theelectrode (1-10 ul Loctite Cyanoacrylate; Henkel Corp., Rocky Hill,Conn.). For cell culture, we used plasma-treated polystyrene fromCorning (Corning, N.Y.) for all experiments involving SY5Y cells and allexperiments involving MCC were performed with poly(lysine) (PDL)-coatedcultureware (BD Biosciences, San Jose, Calif.) or dishes and electrodescoated with 1 mg/ml PDL (Sigma-Aldrich) in PBS for 2-12 h (then rinsedin PBS prior to cell exposure). For experiments in which PEDOT waspolymerized around the living cells, the neural cells were cultured onthe electrode for 24-48 h prior to electrochemical polymerizationprocess.

Electrochemical polymerization and removal of cells from PEDOT: Theelectrode was placed in an electrically-connected reservoir containingthe aqueous monomer solution (for these studies: 0.01 M EDOT and 0.02Mpoly-anionic dopant poly(sodium styrene sulfonate) (PSS) in phosphatebuffered saline (PBS; Hyclone Media, Logan, UT). Galvanostatic current(0.5-10 uA/mm²) was applied to the electrode and the monomer solutionusing an AutoLab PGstat12 Potentiostat/Galvanostat (EcoChemie, TheNetherlands) for 0.5-10 minutes depending on the geometry of theelectrode and the desired thickness of the polymer film. For studies oncell-templated PEDOT substrates, cells are cultured on electrodes, PEDOTwas electrochemically deposited around the cells, immediately followingpolymerization the cells were removed by exposure to 100 mMtrypsin-versene (Hyclone) at 37° C. for 2 h followed by mechanicaldisruption.

Microscopy. We used several different types of microscopy tocharacterize the interactions between electrodes and neural cells. 1)Optical microscopy: Nikon Optiphot POL with a Spot RT digital camera; 2)Phase contrast/fluorescence microscopy: Nikon T2000 invertedlight/fluorescence microscope with Hg arc lamp, Hamamatsu CCD 16 bitcamera with Simple PCI imaging software (courtesy of Takayama lab);upright Olympus BX-51 with Hg arc lamp, Olympus CCD camera, and Olympusimaging software (University of Michigan Microscopy and Image AnalysisCore Laboratory, MIL); 3) Scanning Electron Microscopy (SEM) andEnvironmental SEM (ESEM): FEI Quanta 3D Dualbeam Focused Ion Beam(University of Michigan Electron Microbeam Analysis Laboratory, EMAL);4) Atomic Force Microscopy (AFM): Digital Instruments Nanoscope III witha multimode head, tapping mode (EMAL).

Cell Viability Assays: Cell viability was assessed using three assays;the Vybrant Live/Dead Assay (Molecular Probes), the MTT cell viabilityassay (Chemicon, Temecula, Calif.), and immunocytochemistry (ICC) forthe apoptosis associated protease, activated caspase 3 (Cell SignalingTechnologies, Beverly, Mass.). For Vybrant Live/Dead assay cellquantity, size, and type of nuclear staining intensity were assessed byfluorescence microscopy using 3 different dyes, Hoechst 33342 (labelsall cell nuclei but brighter in nuclei of apoptotic cells), YoPro-3(labels apoptotic cells) and propidium iodide (PI) (labels cells withcompromised membrane; apoptotic & necrotic cells).

Immunocytochemistry and Cell Staining: Cells were fixed in 3.7%formaldehyde/PBS at RT for 30 min-1 h. For ICC, non-specific antibodybinding was blocked with 3% BSA/PBS+0.1% Triton X (PBSX), primaryantibodies (activated caspase 3; Cell Signaling Technology, Beverly,Mass.) were diluted 1:100 in blocking buffer and incubated with cellsovernight at 4° C. The next day cells were washed in PBSX, incubatedwith secondary antibody (1:300 in blocking buffer), nuclei werecounterstained with Hoechst/PBS (Molecular Probes/Invitrogen) then cellswere washed, then aqueous mounted with Fluoromount G (Fisher) forimaging. The F-actin cytoskeleton was labeled by Phalloidin-Oregon Green(Molecular Probes) (1:300 in PBSX) for 1 h at RT or overnight at 4° C.For fluorescence microscopy and ESEM, cells were fixed with 4%formaldehyde, maintained in PBS, then washed in water prior to imaging.For SEM cells were fixed using 1% gluteraldehyde, washed in water, thendehydrated in ascending ethanols (50%, 75%, 95%, 100%; 10 min each) thendried overnight in Peldri II or hexamethyidisilazane (HMDS) (Ted Pella,Redding, Calif.).

Electrical Properties Analysis: Electrical testing of electrodes wasperformed before and after PEDOT deposition using the AutoLab PGstat anda 3 electrode system with PBS (pH 7.0) as the electrolyte, a platinumwire as the counter electrode (CE), a saturated Ag/AgCI calomelelectrode (SCE) as the reference electrode (RE). The electrode itselfwas the stimulating/working electrode (SE/WE). Electrochemical ImpedanceSpectroscopy (EIS) was used to assess the response to alternatingcurrent (AC) over a range of frequencies (1-100,000 Hz), paying closeattention to the behavior at 100-1000 Hz, frequencies typicallyassociated with detecting neural activity with microelectrodes. CyclicVoltammetry (CV) was used to determine the charge capacity of theelectrodes. The voltage was cycled from −1 to +1 V or −0.9 to 0.5 V vs.SCE at a rate of (0.1 V/s) while the current was measured.

Equivalent Circuit Modeling: ZSimpWin (EChem Software, Ann Arbor, Mich.)was used to develop a circuit model from the EIS data. Data was importedfrom the AutoLab PGstat software, Frequency Response Analyzer (FRA). Themodeling process was iterative, using the Chi-Square (X²) value for theentire model and the percent error values for each circuit component todetermine the fit of a given model to the experimental data. Componentswere chosen using theories from electrochemical cell studies and usingthe Boukamp suggestion that each component addition should reduce the X²value by one order of magnitude. Circuit models are presented using theBoukamp representation. The X² value was calculated according to thefollowing algorithm:

-   Experimental Data Point [ω_(i), a_(i), b_(i) ]-   Parameters Associated with ρ=(ρ₁, ρ₂ . . . ρ_(m))-   Model [ω_(i), z_(i)′(ω_(i),ρ), Z_(i)″(ω_(i), ρ)]-   Calculated Point [ω_(i), W_(i)′, W_(i)″]-   Weighing Factors

$\chi^{2} = {\sum\limits_{i = 1}^{n}\lbrack {{W_{i}^{\prime}( {{Z_{i}^{\prime}( {\omega_{i},\rho} )} - a_{i}} )}^{2} + {W_{i}^{''}( {{Z_{i}^{''}( {\omega_{i},\rho} )} - b_{i}} )}^{2}} \rbrack}$

-   Chi-Square (X²) Value W₁′=W_(i)″=1.0/(a_(i) ²+b_(i) ²)    The X² value was minimized when the experimental data points    correlate with the theoretical data points. This was done by first    calculating the difference between the experimental and calculated    data points. The difference was squared to give larger variances a    greater significance. The differences for all data points were    summed and then divided by a weighing factor. A X² of on the order    of 1×10⁻³ or below was acceptable for a given model.    Results and Discussion

In order to study interactions between the conducting polymer, PEDOT andneurons in vitro, we used two different types of neural cell cultures,mouse primary cortical cultures (MCC) and SH-SY5Y neuroblastoma-derivedcell line (SY5Y) and two types of electrodes, custom-designed, in-housefabricated Au/Pd sputter-coated (Au/Pd; 6 mm electrode diameter) andApplied BioPhysics (Troy, N.Y., USA) ECIS electrodes (ABP; 250 μmelectrode diameter). Cells can be cultured for days to weeks onconducting polymers such as PEDOT and poly(pyrrole) with little or notoxicity. However effects of the monomer on cell viability were notknown. Therefore we first determined the cytotoxicity dose-responsecurve for serial dilutions of the PEDOT monomer, ethylenedioxythiophene(EDOT) and the poly-anionic dopant poly(styrene sulfonate) (PSS, 0.02M).We found that both SY5Y cells and MCC could be exposed to as much as0.01 M EDOT, 0.02M PSS for as long as 72 h while maintaining at least75% cell viability. Therefore since we typically used PEDOTpolymerization procedures of 30 sec-10 min. in duration, we expectedcytotoxicity would be negligible.

Polymerization of PEDOT around living cells. To investigate whetherPEDOT could be polymerized directly in the presence of live neuralcells, we electrochemically deposited PEDOT using 0.5-1 uA/mm²galvanostatic current from a monomer solution containing 0.01 M EDOT,0.02 M PSS in PBS onto electrodes seeded with neural cells. Thisresulted in formation of PEDOT on the electrode, surrounding andembedding the cells. We assessed the morphology and topology of thePEDOT polymerized around the neural cells using optical microscopy andscanning electron microscopy (SEM). After deposition, PEDOT appeared asa dark, opaque substance around the cells and the cells and their nucleiremained intact throughout and following polymerization. Interestingly,PEDOT deposition was prohibited in areas where cells were evidentlystrongly adhered to the substrate. Using SEM, we found that the PEDOT onthe electrode and around the cells displays the fuzzy, nodular surfacetopology that is typical of PEDOT. The polymer also appeared to wraparound the exterior of the cells and their extensions, in some casesgrowing over, engulfing the cell body.

Generation of cell-templated PEDOT coatings. We next adapted thesetechniques to generate conductive polymer substrates with biomimetictopology consisting of cell-shaped holes and imprints on the same scaleas cell surface features. Following polymerization of PEDOT around theneurons, the cells and cell material was removed from the PEDOT matrixusing enzymatic and mechanical disruption. This resulted in a neuralcell-templated, fuzzy PEDOT material with a combination of nanometer andmicrometer scale features. The neural cell-templated polymer topographyincluded neuron-shaped holes and tunnels, crevasses, and caves resultingfrom conductive polymer molded around cell bodies and extended neurites.Using this method, we found evidence of intimate contact at theinterface between the PEDOT matrix and plasma membrane of the cells inwhich the PEDOT (dark substance) revealed nanometer scale tendrils atthe leading edge of a neurite. AFM images provided further details aboutthe topology of the polymer surface, indicating that theneuron-templated features were about 1.5-3 μm in height.

We hypothesized that the biomimetic surface of the cell-templated PEDOTwould be attractive to cells due to its nanometer scale “fuzziness” andthe unique cell-shaped holes and imprints. After new cells were seededon top of the cell-templated PEDOT, we probed for evidence of cellre-population of the cell-shaped holes or of increased adhesion to thecell-templated surface. We found that SY5Y cells cultured on theneuron-templated PEDOT substrate showed a preference for adhering to thecell templated zones over the regions of un-templated PEDOT. A subset ofcells did seem to re-populate the cell-shaped holes of the film, howeverthese cells did not settle down into the exact position as the originalcells used for templating.

Assessment of cellular responses to embedding in PEDOT. To betterunderstand cellular responses to the electrochemical polymerizationprocedure and embedding within the PEDOT matrix, we assessed cellviability, morphology of the cytoskeleton and nuclei, cell adhesion, andcell membrane integrity during minutes to days after polymerization. Forthese experiments, we used polymerization procedures which resulted inPEDOT matrices that did not completely engulf the cells so that cellularactivity and access to nutrients could be retained. We found that thecells did not undergo lytic or necrotic death as evidenced by normalnuclear morphology (Hoechst 33342 staining) after the first 24 hfollowing polymerization. Therefore we assayed the cells for programmedcell death or apoptosis which can occur 24-96 h following the triggeringinsult using the Vybrant Live/Dead assay (Invitrogen) andimmunocytochemistry for activated caspase 3 (Cell SignalingTechnologies), an apoptosis-associated protease. Indeed, starting at 72h following polymerization we began to detect increasing percentages ofapoptosis in cells embedded in the PEDOT matrix as indicated by thepresence of activated caspase 3 in the nuclei. For example, comparisonof percentages of activated caspase 3 (+) cells in MCC at Oh (FIG. 5 c)and 120 h post polymerization. Apoptotic cell counts at 0 h afterpolymerization revealed few if any apoptotic cells however by 120 hafter polymerization, 25% and 33% apoptotic cells were detected in SY5Yand MCC, respectively.

Cells were stained with propidium iodide (PI), a nucleic acid dye thatis impermeable to cells with intact plasma membranes. The PI (+)staining was transient and by 24 h there was no significant differencebetween electrochemically polymerized cells and controls (no currentexposure). The cells were surrounded by a thick, dense PEDOT matrix(dark, opaque substance) that covered most of the neurites leavingexposed only the tallest cell regions near the soma.

Characterization of the electrical properties of PEDOT containing cells.We next characterized the electrical properties of the neuron-templatedPEDOT and PEDOT+live neuron electrode coatings using ElectricalImpedance Spectroscopy (EIS) and Cyclic Voltammetry (CV). Recording ofelectrophysiological signals from electrically active cells such asneurons and cardiac myocytes are typically performed at frequency rangesfrom 0.1-1 kHz with low impedance, sensitive electrodes which providethe highest signal to noise ratio and number of recordable units.Electrode impedance is related to interfacial surface area between theelectrode and electrolyte with impedance decreasing as surface areaincreases. Consistent with previous reports, coating of electrodes withPEDOT results in lowering of electrode impedance 1-2 orders of magnitudeacross frequencies between 0.01-100 Hz. This is evidently due at leastin large part to an increase in effective surface area of the electrodewhich is provided by the fuzzy, nano-porous yet conductive PEDOT matrix.Compared to PEDOT alone, the impedance of the PEDOT+neurons coating isincreased due to the presence of the cells. This is likely associatedwith decreased PEDOT coverage of the electrode surface because the cellsact as a barrier to PEDOT polymerization on some regions of theelectrode. However our unexpected finding that neural cell-templatedPEDOT coatings showed impedance spectra between that of electrodescoated with PEDOT and PEDOT+neural cells suggests that some of theincreased impedance of PEDOT+neurons compared to PEDOT could be due tothe electrically-active nature of the cells which may interfere withsignal transduction between the electrode and the PEDOT. Compare the1000 Hz impedance (Z) of the bare, uncoated electrode (4.4 kOhms) tothat of an electrode seeded with neural cells (2.7 kOhms), thePEDOT-coated electrode (0.2 kOhms), the PEDOT+live neural cellselectrode (1.3 kOhms), and the neural cell-templated PEDOT electrode(0.7 kOhms).

The phase plot of the impedance spectroscopy reveals phase angles of75-85° for the bare and neural cell-seeded ABP electrodes at frequenciesof <10 kHz indicating that the electrode is primarily functioning as acapacitor. Coating with PEDOT dramatically drops the phase angle to <20°making the electrode more resistive as opposed to capacitive atfrequencies above 0.1 kHz. However the presence of neural cells withinthe PEDOT matrix tempers this response attenuating the decrease in phaseangle so that it does not become primarily resistive until >10 kHzfrequencies. This is likely due to complex interactions between theneural cell membranes which inherently have both resistive andcapacitive components (usually represented by RC circuit withdepolarization resistance R and membrane charge storage capacity C andthe unique microstructure of the PEDOT matrix that forms when PEDOT ispolymerized in the presence of live cells.

To better understand how the PEDOT+neuron and neuron-templated PEDOTcoatings related to PEDOT coatings in terms of their ability to decreasethe electrical impedance of an electrode, we compared them to twosimilar PEDOT coatings that we characterized in previous publications.PEDOT+neuron and neuron-templated PEDOT coatings were compared to aPEDOT coating comprising of an EDOT monomer solution containing the samepoly-anionic dopant, PSS used in the present studies as well as a PEDOTcoating templated with 485 nm poly(styrene) spheres using a methodsimilar to the methods presented here for preparing cell-templatedPEDOT. Because we have used a variety of electrode types and geometriesin our publications, for comparison purposes the data were normalizedfor electrode surface area (Z*A=Ohms*m²) and the 1 kHz impedance valueswere graphed as a function of deposition charge density (C/A=C/m²).

We have also performed equivalent circuit modeling to better understandhow increasing complexity, microporosity, and non-uniformity of PEDOTcoatings can dramatically affect resistivity of the PEDOT coatings. Atypical bare electrode can be represented by RS(T(R_(T)Q)) where R_(s)is solution resistance, T is a diffusion-related finite Warburg element(constant phase element Q_(n)=0.5), R_(T) is charge transfer resistanceat the electrode-electrolyte interface, and Q is a constant phaseelement representing the porosity and interfacial capacitance of theelectrolyte-electrode interface. Previously, equivalent circuit modelsfor PEDOT-coated electrodes have been defined as R(C(R_(T)Q_(n)=0.5)) inwhich the T of the bare electrode is substituted for a C (capacitor) dueto the diffusion of ions at the polymer surface and current conductionthrough the polymer that is more capacitative than for the bareelectrode. Interestingly, modeling calculations for PEDOT,PEDOT+neurons, and neuron-templated PEDOT coatings on ABP electrodesindicated that the PEDOT matrix was best represented by a constant phaseelement Q_(n)=0.97 for PEDOT alone, Q_(n)=0.88 for PEDOT+neurons, andQ_(n)=0.72 for neuron-templated PEDOT (see Table 2). This decreasingtrend represents an increase in the surface porosity of the PEDOT whichcan be corroborated by qualitative analysis of PEDOT, PEDOT+neurons, andneuron-templated PEDOT which indicates that neuron-templated PEDOT hasthe highest gross porosity due to the presence of cell-shaped holes inthe PEDOT matrix.

The presence of neural cells in the PEDOT matrix contributed an RCelement typical of neural cell membranes that was in parallel with theC(RQ) of the PEDOT resulting in [R_(s)(C(R_(T)Q)(RC))]. Interestinglythe same model could be applied to neural cell-templated PEDOT yet inthis case the additional RC was contributed by the capacitative gapsleft behind after removal of cells from the PEDOT matrix rather than bythe effects of the cell membranes. Despite having the same model, thevalues for both the resistor and the capacitor in the RC element arehigher for the PEDOT+neuron coating (R=3.18×10⁻³ Ohms cm², C=2.46×10⁻¹F/cm²) as compared to the neuron-templated PEDOT coating (R=7.14*10⁻³Ohms cm², C=7.66×10⁻¹ F/cm²). This increase in resistivity andcapacitance manifested in an increase in charge transfer capacity thatcan explain why neuron-templated PEDOT is more conductive than thePEDOT+live neuron matrix.

Cyclic voltammetry was used to assess the charge transfer capacity ofthe PEDOT, PEDOT+neurons, and neuron-templated PEDOT coatings on Au/Pdelectrodes. The dramatic increase in charge capacity (area under CVcurve) for PEDOT and neuron-templated PEDOT-coated electrodes ascompared to the bare Au/P electrode is consistent with PEDOT coatings.CV spectra show the intrinsic redox reaction of the electrode materialas the potential of the electrode bias is cycled from negative topositive and back. This propels ion exchange between the electrode andthe electrolyte moving mobile charge carriers in and out of the PEDOTmatrix. This voltage bias switching process can be repeatedly applied toPEDOT-coatings with little or no degradation of the electrical orphysical stability of the film, making PEDOT-coated electrodes idealcandidates for biosensing and drug-releasing biomaterials applications.The charge capacity for the PEDOT+neuron electrode coating is alsogreatly increased over the bare Au/Pd electrode seeded with neural cellsbut does not reach the level of that seen for PEDOT and has a distinctlydifferent shape.

Interactions between neural cell cultures and the conducting polymerPEDOT are advantageous for the development of electrically conductivebiomaterials intended for contact with electrically-active tissues suchas the brain and heart. PEDOT was electrochemically polymerized directlyin the presence of neural cells seeded on electrodes resulting in theformation of a conducting polymer matrix around and onto adhered cells.SEM and optical imaging suggested that polymerization from a monomersolution enabled the polymer to deposit at the cell-electrode interface,apparently using the cells, cell membranes, and extracellular matrix(ECM) as scaffolds for polymerization.

Electrical characterization of the PEDOT matrix containing live neuralcells suggested a relationship between the electrode and neural cellsthat is distinct from a more typical configuration used for electricallyinterfacing neurons in which neural cells are cultured on or near metalelectrodes. Intimate interactions between the conducting polymer and theneuronal membrane were revealed as PEDOT covered delicate filopodia andneurites. This unique cell-polymer-electrode interface can be an idealcandidate material for the development of a new generation of biosensorsand “smart” bioelectrodes. The incorporation of electrically-responsive,electrode-adherent cells into a conducting polymer matrix provides foran additional opportunity to exploit both the biochemical andelectrochemical qualities of the incorporated cells for sensingpurposes.

PEDOT polymerized around cells cultured on electrodes also indicatedthat the process of electropolymerization around living cells is a novelmethod for capturing and immobilizing cells in a fixed, conductivematrix. Trapping cells on an electrode site in PEDOT can simplifymulti-electrode array (MEA)-based electrophysiological studies ofsignaling in neural networks which is currently made difficult bymigration of neurons off electrode sites. Electropolymerization of cellson an electrode can facilitate imaging of cells using Atomic ForceMicroscopy (AFM) and Scanning Tunneling Microscopy (STM) which requireconductive substrates and/or immobilized targets. We also noted thatPEDOT polymerized around cells cultured on electrodes is a novel methodfor revealing a “negative” image of the morphology of the cell-substrateadhesions due to the manner in which PEDOT is deposited around theexterior of the cells. This can provide an alternative to other methodsfor visualizing cell-substrate interactions such as immunocytochemistryand Total Internal Reflection Fluorescence (TIRF) microscopy.

We next generated neuron-templated PEDOT coatings by removing the cellsand cell material from the PEDOT matrix after polymerization around thecells. We hypothesized that a cell-templated surface would becytomimetic, probably biocompatible and possibly cell-attractive. Indeedthe cell-defined PEDOT matrix provided surface features on the cell andneurite length-scale and included tunnels, troughs, crevasses, and cavesresulting from PEDOT molded around extended neurites and variouscellular processes. Our in vitro findings presented herein indicatesthat when implanted in tissue, this cell-templated polymer surface canencourage cells in the host tissue to re-populate the cell-shaped holesand send processes into the tunnels and crevasses. This would providefor very intimate contact between cells and the conductive polymermaking possible continuous electrical contact between the electrode andthe tissue. Variation in cell removal techniques can provide anopportunity for spatially-localized biochemical control of interactionsbetween target cells and the electrode at the cellular and subcellularlength-scale. When coupled with the mechanical control provided by thecytomimetic topology, tailoring of the biochemistry of thecell-templated surface could make possible precise manipulation andtracking of neurite guidance, growth, and signal transduction.

Consistent with other conducting polymer electrode coatings, thecell-templated PEDOT and PEDOT+neuron coatings described hereindemonstrates the ability to enhance electrode functionality as indicatedby decreased electrical impedance of at least 1 order of magnitude at 1kHz and charge capacity increases of 2-4× the bare electrode. Hence,paired with their biomimetic properties, these novel electrode coatingsare excellent candidate materials for improving the electrode-tissueinterface.

Example 3 Bioelectrodes Comprising Cells Contained Within a HydrogelScaffold

The bioelectrode comprises living, active cells, immobilized in a 3Dhydrogel scaffold with conductive polymer networks deposited through thegel and around cells. The conductive polymer allows the bioelectrode torelay electric or electronic signals to and from other devices forelectrical communication. The cells suspended within the hydrogel can bemonitored directly with the electrode, and can also be used tobiochemically or electrically interact with the desired communicationsource.

Methods and Materials: Tissue/cell culture: Cells are harvested asdescribed in Example 1. For neuronal cells as prepared according toExample 1, are plated on poly(lysine) coated cultureware. Glial cells inthe culture can be limited by not including FBS in the media used formedia changes. Cells are grown according to Example 1. Prior toimmobilization in the hydrogel scaffold, cells cultured in dishes areenzymatically removed from the substrate by incubation with Trypsin-EDTA0.25% (Gibco/lnvitrogen) at 37° C. for 10-15 minutes. The cells andmedia are the centrifuged at 1000 RPM for 2 minutes and the supernatantis discarded. The remaining cell pellet is resuspended and triturated todissociate cells in enough media for a concentration of 10⁵-10⁷cells/ml.

In certain embodiments, alginate and poly(vinyl alcohol) (PVA) hydrogelscan be used however, the method can be adapted to a number ofbiocompatible hydrogels or chemically-functionalized hydrogels withnon-toxic crosslinking. Alginate hydrogels are made from high G, mediumviscosity alginate powder (Sigma Aldrich, St. Louis, Mo.) dissolved inPBS (1-6% (w/v)) and then filter sterilized using 0.45μm syringe filters(Fisher Scientific, Hampton, N.H.). The alginate:PBS solution is thenthoroughly mixed with enough of the cell solution for a cellularconcentration of 5×10⁴-5×10⁶ cells/ml and an alginate concentration of0.5-3% (w/v). Crosslinking is achieved by addition of a sterilizedsource of divalent ions, such as Ca²⁺ or Mg²⁺. Thin (5μm-2 mm) electrodecoatings of hydrogel containing living cells are applied by dippingelectrodes or wires into the hydrogel-cell solution and then bysubmerging the electrode in a 2% (w/v) CaCI₂ (Sigma-Aldrich) solution indeionized water which has been sterilized using a 0.22 μm syringe filter(Fisher Scientific). Repeated hydrogel applications and crosslinkingscan be used to create thicker coatings. Larger (up to 10 cm³) bulkhydrogels are made by thoroughly mixing the hydrogel-cell aqueoussolution with a filter sterilized 4% (w/v) CaSO₄ solution in deionizedwater in a molar ratio of 0.18. The gel is then injected into asterilized mold or receptacle of choice. Hydrogel scaffolds can betemporarily stored in PBS or HBSS during fabrication of thehydrogel-space filling bioelectrode.

Electrochemical polymerization in the presence of the living cells: Thecell-seeded hydrogel-coated electrode can be placed in anelectrically-connected reservoir containing a saline solution such asPBS or HBSS that contains the desired monomer with dopants and/orbiomolecules. For bulk hydrogels, an electrode or microwire is insertedinto the hydrogel. Galvanostatic current is applied to the electrode andthe monomer solution using an AutoLab Potentiostat/Galvanostat(EcoChemie, The Netherlands) or some similar instrument capable ofdelivering direct current (DC) at 1-10 μA/mm² for 0.5-120 minutesdepending on the desired thickness of the polymer film. Electrochemicaloxidation/reduction of the monomer results in the formation ofconducting polymer films and networks through the hydrogel network andaround the cells, thus embedding and immobilizing them in a 3Dconductive polymer hydrogel scaffold.

Assessment of cell viability: Viability of cells in contact with theconductive polymer electrode networks can only be assessed as describedin Examples 1 and 2.

Characterization/measurement of functionality & effectiveness:Characterization of surface morphology: Assessment of the electricalproperties: The combination of the micrometer and nanometer scalesurface roughness of the conducting polymer film/network and thecell-templated pores and tubes that result from electrochemicallypolymerizing in the presence of the living cells can manifest in anincrease in effective surface area of the electrode and thussignificantly decreases the electrical impedance while increasing thecharge capacity of the electrode. To assess these changes, we measurethe electrical properties of the 3-dimensional electrode network byperforming Electrical Impedance Spectroscopy (EIS) and CyclicVoltammetry (CV). We use the Brinkmann Autolab system connected to aDell computer to perform these measurements. A solution of 0.1 M PBS (pH7.0) is used as the electrolyte in a three-electrode cell. A platinumfoil is used as the counter electrode and a saturated calomel electrodeis used as the reference electrode. The conductive polymer electrodenetwork is connected to (and becomes) the working electrode.

For EIS, an AC sinusoidal signal of 5 mV amplitude is used and the DCpotential set to 0. The values of the impedance are determined at fivediscrete frequencies per decade over the range of 10⁵-10 Hz. The realand imaginary components of the impedance are measured as a function offrequency and plotted in various format (amplitude vs. frequency, phaseangle vs. frequency, real part vs. imaginary part) for analysis. For CV,the three-electrode cell setup is the same as the one used for EIS. Ascan rate of 10 mV/s will be used and the potential on the workingelectrode will be swept between-1.0 to 1.0 V vs. SCE. This limit is wideenough to include the reversible redox reaction and narrow enough toavoid over-oxidation and remain in the water window.

Example 4 In situ Polymerized Electrode Networks

In this embodiment, the conductive polymer is a diffuse network ofmolecularly-thin and nanometer scale conductive polymer fibrils that isgrown in situ through interstitial spaces in tissues and within theextracellular matrix between cells. To fabricate the bioelectrode withina tissue, it can be necessary to have an electrode substrate on animplantable biomedical device from which the polymerization of theconducting monomer is achieved by the delivery of electrical currentthrough the electrode site. For polymerization of the conductivepolymer, the tissue near the electrode site must be saturated in thenon-toxic monomer solution which can be accomplished by deliveringmonomer via microfluidic channels in the biomedical device or byinjection.

When electrical current is delivered in the presence of the monomersolution, the polymer electrode first deposits on the electrode siteitself and then grows off the electrode site, following the electricfield generated by the electrode. The result is creation of anelectrically-connected diffuse network of thin polymer fibers and chainswoven around and between cells, effectively innervating the tissue andintimately contacting the plasma membranes of cells in the 3D space ofthe living tissue. The diffuse polymer network electrode is fullyintegrated within the living tissue and it maintains electricalintegrity and stability as it moves with the tissues, eliminatingmicromotion-associated tissue damage that is often seen with physicallytethered or stiffer electrodes that are not well-integrated at thetissue-device interface.

In certain embodiments, the conductive polymers can be polymerizedwithin a variety of tissues including, but not limited to, epithelialtissue, dermis, cardiac muscle, and brain. Due to its ability to growwithin & through nanometer thin spaces between cells for lengths of atleast 500 um to 1 mm from the electrode site, this type of 3-dimensionalelectrode network can penetrate and bypass fibrous scar encapsulationsand congregations of immune cells that often form around implantedelectrodes. This makes possible the establishment of functionallong-term electrical communication between implanted biomedical devicesand the healthy target cells tissue despite encapsulation of the devicein high electrical impedance and the presence of signal blocking scartissue.

Synthesis of Diffuse Polymer Electrode Networks: An electrode or abiomedical device with 1 or more electrode sites is inserted into thetarget tissue. The electrode must be electrically connected (throughelectrical wires or some sort of telemetry) to an instrument and/orcomputer that can deliver electrical stimulation to the electrodes ofthe implanted device. For polymerization of the conductive polymer, thetissue near the electrode site must be saturated in the non-toxicmonomer solution which can be accomplished by delivering monomer viamicrofluidic channels of the biomedical device or separately byinjection or infusion.

Polymerization in tissue for acute in vitro testing: The tissue (e.g.brain, heart, skin, muscle, etc.) in which the Diffuse Polymer ElectrodeNetwork is to be polymerized is dissected from an euthanized (CO₂overdose) adult Swiss-Webster mouse and immediately submerged in icecold monomer solution for 10-30 minutes at 4° C. The monomer solution isa saline (PBS or HBSS) solution containing 0.01M 3,4 ethylenedioxythiophene (EDOT), 0.25 mg/ml of the ionic dopant poly(styrenesulfonate) (PSS). A variety of other dopants and biomolecules can alsobe included in the monomer solution. After incubation in the monomersolution, the tissue can be placed in an electrically-connectedreservoir filled with chilled monomer solution and a 75 um diameter gold(Au; Teflon-coated) microwire (Ted Pella, Redding, Calif.) electrode isinserted into the desired position in the tissue. Galvanostatic currentis then applied to the electrode and the monomer solution using anAutoLab Potentiostat/Galvanostat (EcoChemie, The Netherlands) or somesimilar instrument capable of delivering direct current (DC) at 1-10μA/mm² that is connected to a computer and electrical analysis software.The polymerization procedure is run for 15 min-4 h at room temperature(RT). Electrochemical oxidation/reduction of the monomer results in theformation molecularly-thin and nanometer scale conducting polymertendrils and networks around the cells in the tissue within theinterstitial spaces. After the polymerization procedure, the tissue(with implanted electrode) can be fixed by submersion in either 4%paraformaldehyde or 2.5% glutaraldehyde (both diluted in PBS) overnight(ON) at 4° C. The next day the tissue is washed in PBS then prepared fortissue sectioning.

Polymerization in organotypic brain slice cultures for chronic testing:Male rats or mice (5-10 days old) are deeply anesthetized by isofluoraneexposure then rapidly decapitated. The brain is removed and placed in anice-cold dissection medium consisting of Hanks buffer with 25 mM HEPESand 6% glucose. The hippocampus and neocortex are dissected out andsliced transversly at 400 um thickness. Slices are placed on planarmicroelectrode arrays (MEAs from MultiChannel Systems Reutlingen,Germany) in 35 mm poly(lysine)-coated tissue culture dishes (BDBiosciences, San Jose, Calif.) or a 75 um diameter gold (Au,Teflon-coated) microwire (Ted Pella) is inserted into the slice and theslice is placed on a semi-porous membrane (0.4 um, Millipore, Billerica,Mass.). Slices with electrodes are cultured in growth media containing50% MEM, 25% horse serum, 25% Hanks buffer, 20 mM HEPES, 1 mM glutamine,and 5 mg/ml glucose at 5% CO2 at 37° C. Cultures are maintained for 3-21days in vitro with media changes every 2-3 days prior to or during usein experiments. 3-7 days after culturing, the tissues are submerged inmedia containing monomer (0.01M EDOT) and allowed to incubate for 1-4 hat 37° C. Next the electrode in contact with the tissue slice iselectrically connected to an AutoLab Potentiostat/Galvanostat(EcoChemie, The Netherlands) or some similar instrument capable ofdelivering direct current (DC) at 1-10 μA/mm² that is connected to acomputer and electrical analysis software. Galvanostatic current is thenapplied to the electrode in the tissue for 15 min-1 h under in vivoconditions (5% CO2 at 37° C.) to maintain cell viability.Electrochemical oxidation/reduction of the monomer results in theformation molecularly-thin and nanometer scale conducting polymertendrils and networks around the cells in the tissue within theinterstitial spaces. After the polymerization procedure the tissue isplaced back in the incubator for a chosen time course followingpolymerization then at the end of the experiment, the tissue (andassociated electrode) is fixed by submersion in either 4%paraformaldehyde or 2.5% glutaraldehyde (both diluted in PBS) overnight(ON) at 4° C. The next day the tissue is washed in PBS then prepared fortissue sectioning.

Diffuse Polymer Electrode Network characterization/ measurement offunctionality & effectiveness: Characterization of Diffuse PolymerElectrode morphology: Diffuse Polymer Electrodes can be synthesizeddynamically, in real-time from an electrode implanted in living tissue.Diffuse polymer electrode networks are another example of a3-dimensional polymer electrode network. Tissues containing the diffusepolymer electrode are usually too thick for imaging by availablemicroscopy methods, the tissue must be sectioned. This can beaccomplished in one of several ways depending on the type of tissuestaining and microscopy to be performed: 1) fixed or un-fixed tissue iswrapped in aluminum foil then flash frozen in liquid nitrogen or dry-icecooled isopropanol then embedded in Tissue-Tek O.C.T Compound (ElectronMicroscopy Sciences, Hatfield, Pa.) and microsectioned (4-20 um slices)by Cryostat; 2) fixed tissue is dehydrated, xylene processed, embeddedin paraffin, then microsectioned (4-12 um slices) by microtome; 3)un-fixed tissue is embedded in 10% gelatin (<50° C.) then sectioned(20-500 um) by vibratome; 4) fixed tissue is embedded in 10% gelatin,the tissue+gelatin complex is fixed ON at 4° C., then sectioned (20-500um) by a vibratome.

Once sectioned, the diffuse polymer electrode networks are evaluatedmicroscopically to assess cell viability, cell morphology, andintegrity/quality of the fully integrated hybrid tissue-conductingpolymer network using optical and fluorescence microscopy. In additionthe surface topography/features are explored using AFM in tapping modein an aqueous environment and as well as environmental scanning electronmicroscopy (ESEM) which is performed in a very low vacuum on a chilledstage (Peltier stage) with 50-70% humidity in the chamber. Opticalmicroscopy is conducted with a Nikon Optiphot POL, having the capabilityfor both reflected and transmitted light observations. Images areacquired with a Spot RT digital camera running on a computer.

For fluorescent microscopy we use Olympus IMT-2 upright light microscopewith Hoffman modulation contrast and a Leica DMIRB fluorescent invertedmicroscope both with mercury arc lamps for UV light, Olympus CCDcameras, and accompanying Olympus digital imaging software running onDell PC computers. In addition for thick tissue sections (>20 um) we usea Zeiss LSM 510 confocal microscope mounted on a Zeiss Axiovert 100Minverted microscope with UV, Argon, and 2 green HeNe lasers that deliverup to four images with transmitted light images and the accompanyingZeis META digital image analysis software that is run on a Dell PC.Information about the surface and the microstructure of the in situpolymerized bioelectrodes will be obtained using the FEI Quanta 200 3DFocused Ion Beam Workstation and Environmental Scanning ElectronMicroscope and a Philips CM-100 transmission electron microscope (TEM)equipped with an automated compustage and Kodak 1.6 Megaplus highresolution digital camera.

Assessment of the diffuse polymer electrode network electricalproperties: The network of molecularly-thin and nanometer scaleconducting polymer fibrils that results from electrochemicallypolymerizing on the scaffold of the living tissue manifests in a largeincrease in effective surface area of the electrode and thussignificantly decreases the electrical impedance while increasing thecharge capacity of the electrode. To assess these changes, we measurethe electrical properties of the diffuse polymer electrode networks byperforming Electrical Impedance Spectroscopy (EIS) and CyclicVoltammetry (CV) as described in Examples 1 and 2.

Example 5 All Polymer Electrodes

In certain embodiments, Polymer wires/electrodes are non-metallic,non-ceramic, and do not contain metalloids (e.g. Silicon) or alloys.Polymer electrodes are comprised of a conducting polymer or combinationsof conducting polymers and non-conductive polymers or hydrogelsjuxtaposed in specific configurations resulting in an electrode leadthat can be used in place of “normal” metal electrodes or wires. In someembodiments, the polymer electrode may also contain carbon or carbonnanotubes. All polymer wires/electrodes offer at least 2 majoradvantages over more traditional metal electrodes. Polymer electrodescan be used in any situation in which it would be unfavorable,dangerous, or impossible to use metal such as in the presence of amagnetic field (e.g. MRI scans of individuals with implanted devicesthat contain metal electrodes devices or bioprosthetics). Secondly,polymer electrodes can be created several ways from a diversity ofsubstrates and materials and are highly adaptable and can be readilytailored for specific, diverse applications from chemical sensing totissue engineering to the next generation of laboratory and scientifictesting/analysis equipment. In addition the polymers comprising thesepolymer electrodes can be prepared to contain or release bioactiveagents which can facilitate electrode (or device) function andcommunication/integration at the interface between the electrode and theelectrolyte.

Applications in which electrodes must be deployed in aqueousenvironments will likely benefit most from use of all polymer wires.This is due in part to the fact that the function of metal wires isoften compromised in aqueous environments as well as the fact that thefunction of all polymer electrodes is in part dependent on and can beenhanced by interaction with electrolytes in the aqueous environments.In addition specific ionic interactions between the polymer electrodeand the electrolyte can be exploited to facilitate the function of thedevice to which the polymer electrode is communicating-similarflexibility is not inherent to more traditional metal electrodes. Fourpossible designs for all polymer electrodes are presented in theaccompanying figures.

Methods and Materials. A non-degradable tubular polymer container 0.5cm-10 cm in length having a diameter of 1-5 cm is sterilized. Polymericmicro/nano fibers, 0.5-9 cm in length are aligned inside the container.Solution of EDOT monomer and appropriate dopant e.g. poly(styrenesulfonate) is placed in the container and the fibers are connected via alead to a source of current. Polymerization of the conducting monomer iscarried out as described in Examples 1 & 2. Excess reagents are removedleaving behind and the conductive polymer coated fibers. An electricalconnection is affixed at one end of the container. Hydrogen canoptionally be poured into the container to coat and surround theconductive micro/nano fibers. The hydrogel can optionally contain livingcells, for example stem cells or biodegradable micro/nano particlescontaining bioactive agents, for example drugs, pharmaceuticals, enzymesgrowth factors and the like. The container can be implanted into thebioelectrode implantation site and connected to a power source via apolymer lead with a metal or conducting polymer.

Polymer Electrode characterization/measurement of functionality &effectiveness. To assess whether the all polymer electrodes areelectrically active, functional electrodes, the electrodes are tested asdescribed in Examples 1-3.

All polymer electrodes can be implemented as special leads that can beattached to devices via traditional metal wires. Therefore it isnecessary to have a device and its various leads/wires available forconnection to the all polymer electrodes/leads. We expect that adiversity of devices will be compatible with these all polymerelectrodes/leads; bio/ion/chemical sensors, “lab-on-chip” devices,implanted biomedical devices, and bioprosthetics components.

Example 6 Diagnostic and “Lab-On Chip” Devices

Conducting polymer-based coatings can be applied to “lab-on-chip”electrodes via electrochemical polymerization of conducting monomersdescribed by various embodiments of the present teachings. Variousbiological components can be incorporated into the conducting polymermatrix during the electrochemical polymerization process. Thesecomponents include but are not limited to antigens, antibodies,receptors, natural or synthetic membranes containing proteins, syntheticmicro or nanoparticles that are coated with antibodies, antigens, orligand-specific surface coatings (e.g. peptides, nucleic acids,chemicals, receptors, proteins), live cells or organisms (e.g. bacteria,viruses), enzymes, synthetic or natural polymers/macromolecules, andmulti-protein complexes.

These biological components are incorporated into the conducting polymermatrix during the electrochemical polymerization process using one ofthe following methods; 1) the agent is added directly to the monomersolution (may also contain ionic dopants and counterions), 2) the agentis deposited on or adhered to the surface of the bare electrode, 3) theagent is injected directly near the electrode sites once theelectrochemical polymerization process is underway. In all of thesesituations, the total surface area of the microelectrodes is calculated,the microelectrodes of the device are bathed in monomer solution, andthen electrical current is applied (0.5-1.5 uA/mm²) for durationstypically ranging from 30 seconds to 30 minutes. The monomer polymerizesaround the biological components entrapping them in a nanoporous polymermatrix (allows ion flow and mass transport/diffusion through polymermatrix) that forms directly and exclusively on top of the microelectrodesites.

Characterization/measurement of functionality & effectiveness ofconducting polymer-based coatings for “lab-on-chip” electrodes: Theelectrical properties of conducting polymer-based coatings on“lab-on-chip” electrodes are assessed using the same methods describedfor other conducting polymer (e.g. PEDOT)-based bioelectrode coatingsdescribed herein (specifically, live cell bioelectrodes, and celltemplated bioelectrodes). The sensing or stimulating functions of thePEDOT-based coating on “lab-on-chip” electrodes can be assessed uniquelywithin the context of each “lab-on-chip” device. For example: a“lab-on-chip” device senses the presence of a specific antigen throughthe use of antibodies embedded within the PEDOT matrix. When thepathogen is present, antigens on the pathogens surface bind theantibodies within the polymer matrix and thus increase the electroderesistance and/or induce an alteration in surface charge of theconducting polymer. Increased resistance and/or a change in polymersurface charge are detected by the underlying electrode which transmitsthis information to the device so that the device can report successfuldetection of the antigen. In this way, the functionality andeffectiveness of the PEDOT-based sensor/coating on the “lab-on-chip”electrodes can be assessed using multiple parameters that depend on boththe electrical properties of the electrode coatings as well as thebioactive/sensing properties of the coatings. It is preferred that thesensitivity and specificity of the pathogen detecting capabilities aremeasured and compared to existing/more traditional methods of detectingthe pathogen of interest.

To best exploit both the unique electrical properties of conductingpolymers and the specific detecting functions provided by the biologicalcomponents incorporated into the conducting polymer matrix, a variety ofelectrochemical analyses will be used to enable detection of bindingevents between the biological components in the conducting polymermatrix and the “target analyte” in the sample solution. This includesbut is not limited to potentiometry, amperometry, cyclic voltammetry,capacitative coupling, and/or electrical impedance spectroscopy. Thespecific electrochemical analysis to be used by a device will depend onwhich type of biological component is present in the conducting polymermatrix and which target analyte in the sample solution is to bedetected. The various electrochemical analysis methods are describedbelow and examples (within the context of a conducting polymer matrix,“lab-on-chip” application) are given for their use.

Potentiometry: In this electrochemical analysis method, voltage orpotential is measured under zero current flow conditions using a2-electrode system, a cathode and an anode. The voltage differencebetween the cathode and the anode is considered the potential of theelectrochemical cell. For lab on chip device applications, theconducting polymer matrix containing the biological component is coatedon the first electrode substrate (usually the cathode). The samplesolution is bathed across the first and second electrodes and the targetanalyte in the sample solution can bind to the biological component inthe conducting polymer matrix. This event elicits a change in surfaceenergy of the conducting polymer matrix and underlying cathode byaltering the species of molecules (and their charges) at the firstelectrode surface as well as by possibly inducing conformational changesin either or both the biological component and the target analyte whichalters charge distribution over the bound/complexed agents. The surfaceenergy change causes a voltage difference between the anode and thecathode. This induces a positive detection response from the lab on chipdevice.

Amperometry: In this electrochemical analysis method, the difference inelectrical current between 2 electrodes is measured while constantvoltage is applied to one of the electrodes (considered the workingelectrode). Like potentiometry, this method can also be used to detectchanges in conducting polymer matrix/electrode surface energy andsimilar concepts can be applied.

Voltammetry (Linear Sweep and Cyclic): A 3-electrode set-up is used forthis electrochemical analysis method, the working electrode, counterelectrode, and reference electrode. The voltage (relative to thereference electrode) is swept at a constant rate from one voltage toanother and the change in electrical current is measured throughout theassay. For linear sweep voltammetry, the voltage is swept from a lowpotential (0.5 V to 5 V) to some higher potential whereas for cyclicvoltammetry (CV) a triangular waveform is used in which the voltage isswept from some negative potential to a positive potential then back tothe negative potential (−1V to +1V to −1V). CV is commonly used tomeasure the redox potentials of chemicals and interfaces in electrolytesolutions. For lab-on-chip device applications, CV can be performed onthe first electrode substrate coated with the conducting polymer matrixcontaining the biological component before and after exposure to thesample solution containing the target analyte. The bound/complexedagents will display a unique CV scan with redox peaks(s) located atdifferent positions than what is characteristic of the conductingpolymer matrix prior to exposure to the sample solution or the bindingevent. Due to its ability to detect redox activity, the CV scan can beused for 2 additional analyses; 1) real-time detection of formation of abinding event(s) between the complimentary molecules at the conductingpolymer matrix and 2) detection of degradation or alterations in theelectrical or physical stability of the conducting polymer matrix.

Electrical impedance spectroscopy (EIS): Similar to voltammetry, a3-electrode set-up is used for EIS. In this method alternating current(AC) is applied at a series of increasing frequencies (Hz) and theimpedance (Z) is recorded. Z is similar to resistance in a DCenvironment but in this case because AC the element that would beessentially equivalent to resistance in Ohms law is Z which isdetermined by the relationship between its 3 components, resistance,capacitance, and inductance. For lab-on-chip device embodiments, theimpedance of the first electrically conductive substrate coated with theconducting polymer containing the biological component is measuredbefore and after exposure to the sample solution. Specific binding ofthe target analyte in the sample solution to the biological component inthe conducting polymer matrix will increase the impedance and alter thephase angle of the impedance, and this will elicit a positive detectionresponse from the device. Like CV, EIS can also be used to measure otheraspects of interaction between the electrode and the solution. Forexample, a specific binding event between the complimentary biologicalcomponent-target analyte will be associated with a distinct EIS profilewhereas non-specific binding of agents in the sample solution to theconducting polymer matrix and electrode will also increase impedance butthis will have a unique pattern that is distinguishable from that of aspecific binding event.

Both potentiometry and amperometry can be simple, one-step analyses thatrequire little programming or battery power which is an advantage.Therefore the electrochemical analysis method selected can be based onwhich biological component is to be used as well as on what kinds ofsample solutions the electrode will be exposed to. For example, alab-on-chip device that employs potentiometry would be preferable for anapplication in which it is desired to detect nucleic acids from alaboratory sample that is otherwise comprised of a saline solution. Incontrast, potentiometry is less selective. For detecting antibodies inblood or serum it would be preferable to design a lab-on-chip devicethat employs CV or EIS electrochemical analyses.

Conducting polymer-based coatings are applied to electrodes on a device,thus it is necessary that the electrode substrate(s) of the device areaccessible for polymerization procedures. In addition, because theconducting polymer coatings provide enhanced electrode sensitivity andcharge transfer capacity it is preferable that both the hardware andsoftware associated with device function are capable of transmitting andinterpreting information coming from the lab on chip device.

The description of the present disclosure is merely exemplary in natureand, thus, variations that do not depart from the gist of the inventionare intended to be within the scope of the invention. Such variationsare not to be regarded as a departure from the spirit and scope of theinvention.

1. A biologically integrated bioelectrode device comprising: (i) a firstelectrically conductive substrate; (ii) a biological component; and(iii) a conductive polymer electrically coupling said first electricallyconductive substrate to said biological component to collectively definea bioelectrode, said bioelectrode transmitting or receiving anelectrical signal between the first electrically conductive substrateand one of said biological component and conductive polymer, saidconductive polymer comprising a surface formed by a template biologicalcomponent, wherein most or all of said template biological component isremoved from said surface, and said biological component is coupled tosaid surface.
 2. The biologically integrated bioelectrode deviceaccording to claim 1, wherein said biological component includes one ormore of a tissue, organic living cell, a cellular constituent orcombinations thereof.
 3. The biologically integrated bioelectrode deviceaccording to claim 2, wherein said organic living cell is selected fromthe group consisting essentially of natural or recombinant eukaryoticcells and prokaryotic cells.
 4. The biologically integrated bioelectrodedevice according to claim 3, wherein said eukaryotic cells is selectedfrom the group consisting essentially of cardiac cells, neural cells,muscle cells, stem cells, stromal cells, hematopoietic cells andcombinations thereof.
 5. The biologically integrated bioelectrode deviceaccording to claim 4, wherein said neural cells comprise neurons.
 6. Thebiologically integrated bioelectrode device according to claim 2,wherein said cellular constituent is selected from the group consistingessentially of a membrane, an organelle, an ion-channel, a lipidbi-layer, a receptor, an enzyme, a protein, an antibody, an antigen, anucleic acid, a carbohydrate, and combinations thereof.
 7. Thebiologically integrated bioelectrode device according to claim 1,wherein said bioelectrode further comprises at least one hydrogel inproximate contact with said conductive polymer.
 8. The biologicallyintegrated bioelectrode device according to claim 7, wherein saidhydrogel further comprises a bioactive substance.
 9. The biologicallyintegrated bioelectrode device according to claim 1, wherein saidconductive polymer is chosen from the group consisting essentially ofcopolymers and homopolymers of EDOT, pyrrole, and their functionalizedderivatives and copolymers, polyanilines, salt of polyaniline,polyacetylenes, polythiophenes, a poly(3,4-ethylenedithiathiophene),polymer blends thereof, hybrid polymer-metal materials and combinationsthereof.
 10. The biologically integrated bioelectrode device accordingto claim 9, wherein the polymers of EDOT and pyrrole is chosen from thegroup consisting essentially of poly(3,4-ethylenedioxythiophene)(PEDOT), poly(pyrrole), their derivatives and combinations thereof. 11.The biologically integrated bioelectrode device according to claim 1,wherein said conductive polymer is polymerized around said biologicalcomponent and said first electrically conductive substrate.
 12. Thebiologically integrated bioelectrode device according to claim 1,wherein said first electrically conductive substrate contains aconductor chosen from the group consisting essentially of gold, silver,platinum, palladium, tungsten, nickel, titanium, indium tin oxide,copper, carbon, carbon black, carbon fiber, carbon paste, graphite,doped silicon, ceramic, conductive polymer, and combinations thereof.13. The biologically integrated bioelectrode device according to claim1, wherein said bioelectrode further comprises one or more dopantsparticipating in polymerization of the conducting polymer andtransmitting or receiving said electrical signal between the firstelectrically conductive substrate and said conductive polymer.
 14. Thebiologically integrated bioelectrode device according to claim 1,further comprising a second electrically conductive substrate coupled tothe biological component.
 15. The biologically integrated bioelectrodedevice according to claim 14, further comprising an electrical source ofpower or current operable to communicate with said first and secondelectrically conductive substrates and the biological component usingone or more electrical signals.
 16. The biologically integratedbioelectrode device according to claim 1, wherein said templatebiological component includes one or more of a tissue, organic livingcell, a cellular constituent or combinations thereof.
 17. Thebiologically integrated bioelectrode device according to claim 16,wherein said organic living cell is selected from the group consistingessentially of natural or recombinant eukaryotic cells and prokaryoticcells.
 18. The biologically integrated bioelectrode device according toclaim 16, wherein said cellular constituent is selected from the groupconsisting essentially of a membrane, an organelle, an ion-channel, alipid bi-layer, a receptor, an enzyme, a protein, an antibody, anantigen, a nucleic acid, a carbohydrate, and combinations thereof. 19.The biologically integrated bioelectrode device according to claim 17,wherein said eukaryotic cells is selected from the group consistingessentially of cardiac cells, neural cells, muscle cells, stem cells,stromal cells, hematopoietic cells and combinations thereof.
 20. Thebiologically integrated bioelectrode device according to claim 19,wherein said neural cells comprise neurons.
 21. A biologicallycompatible electrode comprising: (a) an electrically conductivesubstrate; (b) a coating on at least some portion of at least onesurface of said electrically conductive substrate, the coatingcomprising: a biological component, and (ii) a conductive polymercomprising a surface formed by a template biological component, whereinmost or all of said template biological component is removed from saidsurface, and said biological component is coupled to said surface.
 22. Abiocompatible and biomimetic coating for an electrically conductivesubstrate comprising: (a) an electrically conductive polymer, and (b) abiological component, wherein the electrically conductive polymercomprises a surface formed by a template biological component, whereinmost or all of the template biological component is removed from thesurface, and the biological component is coupled to the surface.
 23. Adevice comprising a biocompatible and biomimetic coating according toclaim 17, wherein said device comprises a microelectrode array, lab onchip device, or target analyte detection device.
 24. A method ofelectrically detecting a transfer of electrical signals between livingcells, comprising the steps: applying a voltage or current across saidfirst electrically conductive substrate of the bioelectrode device ofclaim 1 and a second electrically conductive substrate, thereby inducinga voltage or current across said conductive polymer, said secondelectrically conductive substrate being electrically coupled to saiddevice and a power source; and detecting the transfer of electricalsignals with said bioelectrode device.
 25. The method according to claim24, wherein said biological component is chosen from the groupconsisting essentially of cardiac cells, neural cells and muscle cells.26. The method according to claim 24, wherein the detecting stepcomprises detecting the transfer of electrical signals wherein saidsignal is selected from the group consisting essentially of impedance,resistance, capacitance, inductance, and current.
 27. The methodaccording to claim 24, wherein said conductive polymer is chosen fromthe group consisting essentially of copolymers and homopolymers of EDOT,pyrrole, and their functionalized derivatives and copolymers,polyanilines, salt of polyaniline, polyacetylenes, polythiophenes, apoly(3,4-ethylenedithiathiophene), polymer blends thereof, hybridpolymer-metal materials and combinations thereof.
 28. The methodaccording to claim 27, wherein the copolymers and homopolymers of EDOTand pyrrole is chosen from the group consisting essentially ofpoly(3,4-ethylenedioxythiophene) (PEDOT), poly(pyrrole), andcombinations thereof.
 29. A biologically integrated bioelectrode devicecomprising conducting polymer in proximate contact with a firstelectrically conductive substrate and at least one biological componentformed by the steps of: (a) contacting a template biological componentwith said first electrically conductive substrate, said firstelectrically conductive substrate chosen from the group consistingessentially of metals, ceramics, carbon, conductive polymers andcombinations thereof so as to form a biologically interfaced electrode;(b) immersing said biologically interfaced electrode in a solutioncomprising conducting monomer and dopant; and (c) polymerizing saidmonomer around said biologically interfaced electrode by inserting asecond electrically conductive substrate into said solution of step (b)and applying galvanostatic or potentiostatic current to said first andsecond electrically conductive substrates for a sufficient time to coatthe biologically interfaced electrode with conductive polymer; (d)removing most or all of said template biological component from theconductive polymer to expose a surface formed by the template biologicalcomponent; and (e) contacting the at least one biological component withsaid surface to form the biologically integrated bioelectrode device.